Structured detectors and detector systems for radiation imaging

ABSTRACT

Detector module designs for radiographic imaging include first and second layers of scintillator rods or pixel arrays oriented in first and second directions. The first and second directions are transversely oriented to define a light sharing region between the first and second layers. Encoding features may be disposed in, on or between the first and second layers, and configured to modulate propagation of optical signals therealong or therebetween.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims priority to U.S. Provisional Patent ApplicationNo. 62/385,426, Structured Detectors and Detector Systems for RadiationImaging, filed Sep. 9, 2016, which is incorporated by reference herein,in the entirety and for all purposes.

FIELD

This disclosure describes novel implementations of detectors anddetector systems that can be employed for diagnostic x-ray and gamma raymedical and small animal imaging (diagnostic x-ray radiology includingx-ray area, slit, slot, tomosynthesis, CT, phase imaging,intraoral/extraoral dental and radiation therapy imaging, nuclearmedicine imaging, PET imaging, small animal imaging), as well as inindustrial, Homeland Security and scientific radiation imaging.

BACKGROUND

The combining of imaging modalities to offer increased functionality hasproduced a number of useful imaging systems, particularly in medicaldiagnostic and small animal imaging. For example, Gamma ray PET detectorsystems are frequently sold with x-ray computed tomography (CT) detectorsystems (although the PET and CT detector systems are physicallyseparate and therefore do not share detectors or a common imagingspace). A notable attempt at offering an integrated imaging system (inwhich detectors and the imaging space of the system are shared) was aSPECT-PET (nuclear medicine and PET) imaging system which reduced costsby sharing detectors and the imaging space (the volume in which theobject is imaged). Although these SPECT-PET imaging systems were notwell received commercially due to performance compromises nonethelessthey offered interesting functionality since SPECT and PET images couldbe acquired separately or simultaneously in a shared imaging space(thereby avoiding registration error between separately acquired SPECTand PET images and reducing the total scan time). In addition,simultaneous CT-SPECT systems have been proposed (typically using CZT orCdTe) although issues arise due to generally differing collimation andflux rate requirements. Both shared and stand-alone imaging systemsbenefit from the implementation of enhanced radiation detectors.

SUMMARY

Embodiments of the invention utilize one or more different improvementsin high speed detector electronics along with various detector materialsdeveloped for human and small animal medical diagnostic imagingincluding diagnostic x-ray radiology (such as x-ray area, slit, slot,tomosynthesis, CT, dental and phase imaging), radiation therapy imaging,nuclear medicine imaging and/or PET imaging, as well as high energyphysics, inspection, etc., to develop cost-effective, single purpose andmultipurpose integrated detector systems.

The details of one or more embodiments of the invention are set forth inthe accompanying drawings and description below. Other features,objects, and advantages of the embodiments of the invention will beapparent from the description and drawings, and from the claims. Allpublications, patents and patent applications cited herein are herebyexpressly incorporated by reference herein, for all purposes.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a perspective view of a non-coincidence Compton-PET detectorimaging system.

FIG. 2 is a perspective view of an edge-on silicon detector substrate inwhich shielded readout ASICs are mounted within an etched region alongthe bottom edge of the semiconductor detector substrate.

FIG. 3 is a perspective view of a focused planar detector.

FIG. 4 is a schematic view of a coincidence Compton-PET detector imagingsystem.

FIG. 5 is a perspective view of a non-coincidence CT-Compton-PETdetector imaging system.

FIG. 6 is a perspective view of a minifying scintillating fiber arraycoupled to a 1D photodetector structured detector suitable for PC orlimited PCE CT imaging.

FIG. 7 is a perspective view of a one-dimensional structured molddetector system with quantum dots or semiconductor detector materials.

FIG. 8 is a perspective view of a two-dimensional structured molddetector system with quantum dots or semiconductor detector materials.

FIG. 9 is a perspective view of a multilayer detector system with N=4layers used for CT and/or PET detector imaging.

FIG. 10A is a perspective view of an alternate multilayer detectorsystem with N=3 layers used for CT and/or PET detector imaging.

FIG. 10B is a perspective view of a multilayer CT and/or PET detectorimaging system with a face-on back-end detector layer.

FIG. 10C is a perspective view of a multilayer CT and/or PET detectorimaging system with a face-on back-end detector layer.

FIG. 10D is a perspective view of a multilayer CT and/or PET detectorimaging system with a face-on back-end detector layer.

FIG. 11 is a perspective view of a focused two-dimensional structuredmold detector system with quantum dots or semiconductor detectormaterials.

FIG. 12A is a perspective view of a non-uniform discrete structured 3Dscintillator detector module with crossed top and bottom layers thatimplement different scintillator rod materials with differentdimensions.

FIG. 12B is a perspective view of a uniform discrete structured 3Dscintillator detector module with crossed top and bottom layers thatimplement the same scintillator rod materials with the same dimensionswith the top layer and bottom layer overhanging their common surfaceemployed for light sharing.

FIG. 12C is a perspective view of a discrete rod-structured,pixel-structured 3D scintillator detector module which implements aparallel array of discrete scintillator rods with light sharing andoptical readouts at both ends of the array of discrete scintillator rodscoupled to a second layer comprised of an array of discrete scintillatorpixels.

FIG. 13 is a perspective view of a double-sided, semi-continuousrod-structured 3D scintillator detector module with a crossing angle of90 degrees comprised of a single scintillator sheet into which physicalgaps (cut/sawed/etched) are introduced with respect to the top andbottom surfaces creating arrays of semi-continuous discrete scintillatorrods.

FIG. 14 is a perspective view of a single-sided, semi-continuousrod-structured 3D scintillator detector module implements a singlescintillator sheet (a single-sided semi-continuous structuredscintillator sheet with a rod structure) into which physical gaps areintroduced in order to create an array of semi-continuous discretescintillator rods in the top layer of the scintillator sheet.

FIG. 15A is a perspective view of a discrete structured 3D scintillatordetector module with crossed scintillator rods implementing a discreteintermediate layer.

FIG. 15B is a perspective view of a double-sided, semi-continuousstructured 3D scintillator detector module with crossed scintillatorrods implementing a continuous intermediate layer.

FIG. 15C is a perspective view of a double-sided, semi-continuousstructured 3D scintillator detector module with parallel scintillatorrods implementing a continuous intermediate layer.

FIG. 16 is a perspective view of a focused discrete structured 3Dscintillator detector module with crossed top and bottom cylindricalshell layers of parallel arrays of discrete scintillator rods.

FIG. 17 is a perspective view of a discrete structured 3D scintillatorfiber detector module in which a structured intermediate scintillatorlayer (comprised of multiple scintillator detector blocks is positionedbetween and coupled to crossed planar layers of scintillator fibers.

FIG. 18A is a perspective view of an edge-on, planar n-level (n=3)multispectral CT detector comprised of an array of detector elementswhich implement photodetectors coupled to the side faces ofscintillators which may differ in at least one of different scintillatormaterial, different scintillator depths.

FIG. 18B is a perspective view of an edge-on, planar n-level (n=4)multispectral CT detector comprised of an array of detector elementswhich implement photodetectors coupled to the side faces ofscintillators which may differ in at least one of different scintillatormaterial, different scintillator depths.

FIG. 19A is a perspective view of an edge-on, n/m-level CT-PET detectorfor n=1 and m=1 in which the PET detector implements photodetectorscoupled to the side faces of scintillators.

FIG. 19B is a perspective view of an edge-on, n/m-level CT-PET detectorfor n=2 and m=1 in which the PET detector implements photodetectorscoupled to the side faces of scintillators.

FIG. 19C is a perspective view of an edge-on, n/m-level CT-PET detectorfor n=3 and m=1 in which the PET detector implements photodetectorscoupled to the side faces of scintillators.

FIG. 19D is a perspective view of an edge-on, n/m-level CT-PET detectorfor n=3 and m=2 in which the PET detector implements photodetectorscoupled to the side faces of scintillators.

FIG. 19E is a perspective view of an edge-on, n/m-level CT-PET detectorfor n=2 and m=6 in which the PET detector implements photodetectorscoupled to the side faces of scintillators.

FIG. 19F is a perspective view of an edge-on, n/m-level CT-PET detectorfor n=at least 1 up to 6 and m=6 in which the PET detector implementsphotodetectors coupled to the side faces of scintillators.

FIG. 19G is a perspective view (from an end perspective) of a sharededge-on, n/m-level CT-PET detector (n=2, m=4) in which the PET detectorelements are preceded by two adjacent edge-on, CT detector elements.

FIG. 20A is a perspective view of a discrete 3D scintillator andphotoacoustic PET detector module irradiate face-on in which ascintillator block 350 is coupled to a photodetector 250 and an acousticarray 440.

FIG. 20B is a perspective view of a discrete 3D scintillator andphotoacoustic PET detector module irradiated edge-on in which ascintillator block 350 is coupled to a photodetector 250 and an acousticarray 440.

FIG. 20C is a perspective view of a discrete 3D scintillator andphotoacoustic PET detector module irradiate face-on in which ascintillator block 350 is coupled to a photodetector 250 and an acousticarray 440 at the same interface.

FIG. 20D is a perspective view of a discrete 3D scintillator andphotoacoustic PET detector module irradiated edge-on in which ascintillator block 350 is coupled to a photodetector 250 and an acousticarray 440 at the same interface.

FIG. 21 illustrates a T-PET system in which a heart/chest PET scanner620 and a head/neck PET scanner 630 operate cooperatively tosimultaneously acquire PET cardiac and brain images of a patient 600.

FIG. 22A is a cutaway view of a planar, face-on implementation of amonolithic multilayer straw detector comprised of straw fibers withsquare cross sections and shared walls.

FIG. 22B is a cutaway view of a planar, face-on implementation of amonolithic multilayer multiwire proportional counter detector.

FIG. 23A is a perspective view of a structured 3D semiconductor x-raydetector with an array of electrode holes that is electronically-coupledto an attached substrate incorporating readout circuitry with a powerand communication link for connection to a computer.

FIG. 23B is a perspective view of a structured 3D semiconductor x-raydetector with an array of electrode channels that iselectronically-coupled to an attached substrate incorporating readoutcircuitry with a power and communication link for connection to acomputer.

FIG. 23C is a perspective view of a movable protective cover that slidesonto a protective frame holding a structured 3D semiconductor x-raydetector with an attached substrate incorporating readout circuitryincluding a power and communication link.

FIG. 23D is a perspective view of a movable protective cover in placeover the protective frame holding a structured 3D semiconductor x-raydetector and attached substrate incorporating readout circuitry forminga digital x-ray camera for intraoral dental imaging with a power andcommunication link connected to computer.

FIG. 24A is a perspective view of a flat, small area transparent storagephosphor film on a support plate (a transparent storage phosphor filmplate) that can be mounted within a protective frame.

FIG. 24B is a perspective view of a movable protective layer that canslide onto the protective frame of a transparent storage phosphor filmplate.

FIG. 24C is a perspective view of a movable protective layer that isattached to the protective frame of a transparent storage phosphor filmplate by a hinge mechanism.

DETAILED DESCRIPTION

Compton cameras are frequently implemented as multilayer detectors.Photon-tracking Compton camera designs considered for photon energiesencountered in applications such as nuclear medicine and PET imaginginclude a single layer (a front-end detector) which provides 3D detectorproperties by incorporating a stack of face-on detector planes of thesame material such as low-Z Silicon (Si) or moderate-Z Germanium (Ge),essentially a multilayer detector, and a multilayer (dual-layer)configuration which combines a 2D detector first layer (the front-enddetector) and a 2D detector second layer (the back-end detector). Notethat the spatial resolution of the first layer and the second layer inthe multilayer (dual-layer) detector design need not be the same.Furthermore, spatial resolution of detector elements within a layer neednot be the same (e.g., a detector layer that offers higher spatialresolution in the center or a detector layer in which the pixel sizeincrease with depth). Note that alternative single layer Compton cameras(mono layer Compton cameras) developed for nuclear medicine and PETtypically employ a 3D semiconductor detector with tracking capability.

The dual-layer, front-end/back-end detector configuration typicallyconsists of a face-on, planar, 2D Si (low-Z) or 2D Ge (moderate-Z)front-end detector combined with a face-on, 2D high-Z back-end detector.Thus, these two Compton camera configurations described herein canutilize detector layers of the same material (low-Z and moderate-Z forCompton scattering) or different materials (low-Z for Compton scatteringand high-Z for photoelectric interactions) for the detection of photonsin the diagnostic energy range of medical imaging. Flexibility in theselection of detector materials and configuration (often with differenttemporal and/or energy resolution) is not limited to separate layers,and different detector materials and configurations can be employedwithin a detector layer.

Clearly other choices of materials can be made depending on the photonenergy range or if other types of particles (neutrons, muons, etc.) areto be detected. Compton camera designs (as well as x-ray scanning andCT, SPECT, PET, dental and hand-held probe designs are described invarious U.S. patents and patent applications including: R. S. Nelson andZ. L. Barbaric, U.S. Pat. No. 4,560,882; R. S. Nelson, U.S. Pat. No.4,937,453; R. S. Nelson, U.S. Pat. No. 5,258,145; R. S. and W. B.Nelson, U.S. Pat. No. 6,583,420; R. S. and W. B. Nelson, U.S. Pat. No.7,291,841; R. S. Nelson, U.S. Pat. No. 7,635,848; R. S. and W. B.Nelson, U.S. Pat. No. 8,017,906; R. S. Nelson, U.S. Pat. No. 8,115,174;R. S. Nelson, U.S. Pat. No. 8,115,175; R. S. Nelson, U.S. Pat. No.8,183,533; R. S. and W. B. Nelson, U.S. Pat. No. 9,384,864; R. S. and W.B. Nelson, U.S. patent application Ser. No. 13/199,612, filed Sep. 6,2011 (U.S. Publication No. 2012/0181437); R. S. and W. B. Nelson, U.S.Pat. No. 9,347,893; R. S. and W. B. Nelson, U.S. patent application Ser.No. 14/804,777, filed Jul. 21, 2015 (U.S. Publication No. 2016/0021674);and R. S. and W. B. Nelson, U.S. patent application Ser. No. 14/804,838,filed Jul. 21, 2015 (U.S. Publication No. 2015/0331115), each of whichis incorporated by reference herein, in the entirety and for allpurposes.

This application also relates to the subject matter of U.S. patentapplications Ser. No. 14/804,777 (U.S. Publication No. 2016/0021674) andNo. 14/804,838 (U.S. Publication No. 2015/0331115), each entitledDETECTOR SYSTEMS FOR RADIATION IMAGING and filed Jul. 21, 2015, whichclaim priority as continuations-in-part to U.S. patent application Ser.No. 13/573,981, entitled COMPTON CAMERA DETECTOR SYSTEMS FOR NOVELINTEGRATED COMPTON-PET AND CT-COMPTON-PET RADIATION IMAGING, filed Oct.18, 2012 (U.S. Publication No. 2014/0110592), which claims priority toU.S. Provisional Application No. 61/689,139, entitled COMPTON CAMERADETECTOR SYSTEMS FOR INTEGRATED COMPTON-PET AND CT-COMPTON-PET RADIATIONIMAGING, filed May 31, 2012, and U.S. Provisional Application No.61/690,348, entitled COMPTON CAMERA DETECTOR SYSTEMS FOR NOVELINTEGRATED COMPTON-PET AND CT-COMPTON-PET RADIATION IMAGING, filed Jun.25, 2012, each of which is incorporated by reference herein, in theentirety and for all purposes. This application further relates to thesubject matter of U.S. patent application Ser. No. 13/199,612, entitledHIGH RESOLUTION IMAGING SYSTEM FOR DIGITAL DENTISTRY, filed Sep. 6, 2011(U.S. Publication No. 2012/0181437), which claims priority as acontinuation-in-part to U.S. Pat. No. 9,384,864, also entitled HIGHRESOLUTION IMAGING SYSTEM FOR DIGITAL DENTISTRY, each of which isincorporated by reference herein, in the entirety and for all purposes.

Compton Camera Detector Systems

Compton camera detector systems exploit the Compton scatter interactionand can also exploit photoelectric interactions (and even pairproduction interactions at sufficiently high photon energies). Comptoncamera detector systems include the capability to track theseinteractions in terms of spatial location and energy deposition with atemporal resolution limited by the detector itself and the readoutelectronics.

Typically the interaction information is used to estimate thedirectionality and energy of the photon incident on the Compton cameradetector system whether the photon is an x-ray, a gamma ray, or anannihilation gamma ray. Note that with the addition of collimation suchas (for example) a pin hole or parallel hole collimator, the Comptoncamera can be converted into a nuclear medicine SPECT camera (Gammacamera). Compton camera features such as tracking capability cancontinue to be utilized. This is an example of a dual-use, integratedCompton detector system in which the types of applications arerelatively unchanged but the capabilities of the detector system aremodified (Nelson, U.S. Pat. No. 7,291,841; Nelson, U.S. Pat. No.8,017,906).

The collimation now provides the directionality of an incident gamma rayindependent of directionality determined by applying Compton camerareconstruction algorithms. It will be shown that the integrated Comptondetector system design can be applied to a range of applications(including nuclear medicine). By employing two or more Compton cameradetector systems with electronic coincidence circuitry (used in medicalPET detector systems) coincidence PET imaging can be implemented.

The flexibility of the Compton camera detector system design(incorporating capabilities such as 3D spatial resolution, energyresolution, detection of photons of different energies, a single layerdetector or a multiple layer detector with the same or differentproperties, photon tracking and coincidence capability) allows versatilenon-coincidence Compton-PET and coincidence Compton-PET detector systemsto be implemented. Furthermore, CT capability can be implemented in theCompton camera detector system design, including non-coincidence andcoincidence Compton-PET designs resulting in CT-Compton-PET detectorsystems. A simplification of this design in which the CT detector andthe Compton-PET detector (or just a PET detector) function independentlywill be referred to as a limited CT-Compton-PET detector system.Furthermore, limited implementations of Compton camera detector designscan be employed for dedicated applications such as (but not limited to)CT imaging or PET imaging.

Although applications discussed herein are primarily directed at medicaldiagnostic x-ray and gamma ray radiation detection, in principle theinvention can also be used to detect radiation such as charged particles(alphas, betas, protons, muons, etc.) and neutrons (as well as otherneutral particles) for the applications described. Furthermore, thedetector systems described herein can be combined with or integratedwith other imaging modalities such as MRI scanners, optical scanners,ultrasound scanners, opto-acoustic scanners, microwave scanners, etc. Itshould be understood that the variations of the dual-use detectorsystems (triple-use, etc. detector systems can also implemented)described herein can be employed for simultaneous or non-simultaneousimaging as required by the appropriate application.

The invention provides detector designs that employ one or more layersof detector modules comprised of edge-on or face-on (or tilted)detectors or a combination of edge-on and face-on detectors (as well astilted detectors). Edge-on detectors (and tilted detectors) canincorporate sub-aperture resolution (SAR) capabilities and face-ondetectors can incorporate depth-of-interaction (DOI) capabilities. Oneor more types of detectors can be employed, including: scintillatordetectors, semiconductor detectors, gas detectors (including, but notlimited to, straw arrays, microstrip gas chambers, multiwireproportional counters and crossed strips multilayer proportionalcounters, gas electron multiplier (GEM), micromegas (Micro-MEsh GaseousStructure) and resistive plate chamber (RPC) detectors), low temperature(such as Ge or superconductor) detectors and structured detectors.

Detectors can offer block, 1D, 2D or 3D spatial resolution as well asadequate, fast or very fast temporal resolution (depending on theapplication requirements). Detectors can offer fixed or adjustablepixels sizes which can be uniform or non-uniform (for example,increasing pixel length along the depth dimension as a function of depthto compensate for beam hardening with depth in a CT detector). Theeffective pixel length along a detector column can be synthesized fromthe outputs of one or more uniformly spaced pixels. Parallel or focusedpixel structures can be implemented. Detectors can operate as energyintegrators, photon counters (PCs) and photon counters with energyresolution (PCEs). Possible detector formats include, but are notlimited to, planar (and focused planar) and focused structure (parallelplanes, ring, partial ring as well as focused ring and focused partialring) detector geometries.

The invention provides novel detector designs and systems for enhancedradiation imaging including Compton and nuclear medicine imaging, PETimaging and x-ray CT imaging. The invention also provides integrateddetector systems based on Compton camera designs.

In one aspect, the invention provides integrated non-coincidenceCompton-PET detector imaging systems. In another aspect, the inventionprovides integrated coincidence Compton-PET detector imaging systems. Inyet another aspect, the invention provides limited integratedCT-Compton-PET detector imaging systems. In still another aspect, theinvention provides integrated non-coincidence CT-Compton-PET detectorimaging systems. In another aspect, the invention provides integratedcoincidence CT-Compton-PET detector imaging systems. Since theintegrated nature of these Compton camera detector designimplementations is readily apparent the term “integrated” willfrequently be omitted when referring to them. Therefore “integratednon-coincidence Compton-PET detector imaging systems” will also bereferred to as “non-coincidence Compton-PET detector imaging systems,”etc. In still another aspect, the invention provides variations ofCompton camera detector designs that can be implemented for dedicatedapplications such as (but not limited to) CT imaging or PET imaging.

The invention employs a range of detector types and formats. The use ofgas, scintillator, semiconductor, low temperature (such as Ge andsuperconductor) and structured detectors in edge-on and/or face-ongeometries has been described for both medical and non-medical imagingapplications. Medical imaging applications include diagnostic x-rayimaging (such as single and multiple x-ray tube sources employed withsingle energy or multiple energy implementations of slit scanning, slotscanning, area radiography, (single or multi-layer) flat panel or planarcone beam CT, focused structure ring or partial ring fan beam CT,focused cone beam CT, tomosynthesis, phase (PCI), coherent scatter,radiation therapy and intraoral/extraoral dental imaging), nuclearmedicine imaging (Compton camera, SPECT/gamma camera detector imagingsystems as well as hand held probe detectors) and PET imaging.Non-medical imaging applications include high energy physics, x-ray andgamma ray astronomy, industrial radiography, Home Land Security (HLS)and military applications. Furthermore it has been shown that detectorspatial resolution can be enhanced using sub-aperture resolution (SAR)or depth-of-interaction (DOI) readout techniques with edge-on andface-on detector geometries, respectively.

Detectors may be layered (stacked) and detector modules within a layercan be partially or completely offset from neighboring detector modules.Individual detectors may function as energy integrators, photon counters(PCs) or photon counters with energy resolution (PCEs), depending on theapplication. One or more of these detector types can be employed withina detector imaging system. (Photon counting (PC) is often mixed up withphoton counting with energy resolution (PCE) in the literature. Photoncounting functions as a (one energy bin) single channel analyzer or SCA.Photon counting with energy resolution functions as a multi-channelanalyzer or MCA).

High speed electronics is provided for tracking interactions andanalyzing the readout signals. An electronic communications link isprovided to a computer for data post-processing, storage, and display.One or more tracking capabilities such as examining nearest neighborpixels for effects related to induced signals and charge diffusion,following scattered or characteristic x-ray radiation within a detectorlayer and between detector layers (if there is more than one detectorlayer), following Compton scattered electrons and photoelectrons andmeasuring coincidence events (for example, the detection of pairs ofannihilations photons in PET imaging), etc., can be implemented.Tracking techniques are used in photon counting and spectral x-rayimaging, SPECT, PET, Compton cameras, hand-held radiation detectorprobes, neutron detectors, detectors with SAR or DOI capability and highenergy physics particle detectors.

Various Compton camera implementations incorporate one or multipledetector layers. These detector layers provide suitable 2D or 3D spatialresolution, energy resolution, temporal resolution, stopping andscattering power and tracking capability. Compton camera, nuclearmedicine SPECT/gamma camera and PET detector imaging systems, tracking,x-ray CT and slit and slot scan detectors, hand held probe detectors,edge-on and face-on detectors (with or without SAR or DOI capability),integrating, PC, and PCE detectors, multi-material detectors along withplanar and focused structure detector geometries have been described invarious U.S. patents and patent applications including Nelson et al.,U.S. Pat. No. 4,560,882; U.S. Pat. No. 4,937,453; U.S. Pat. No.5,258,145; U.S. Pat. No. 6,583,420; U.S. Pat. No. 7,291,841; U.S. Pat.No. 7,635,848; U.S. Pat. No. 8,017,906; U.S. Pat. No. 8,115,174; U.S.Pat. No. 8,115,175; U.S. Pat. No. 8,183,533; U.S. patent applicationSer. No. 13/199,612 (U.S. Publication No. 2012/0181437); U.S. Pat. No.9,347,893; U.S. patent application Ser. No. 14/804,777 (U.S. PublicationNo. 2016/0021674) and U.S. patent application Ser. No. 14/804,838, (U.S.Publication No. 2015/0331115), each of which is incorporated byreference herein, in the entirety and for all purposes.

X-ray or gamma ray interactions (in medical imaging applications) can betracked between sufficiently thin detector layers, each with 2D spatialresolution capability. If the depth of a 2D detector layer issufficiently small such that tracking position errors are acceptablethen it effectively functions as a restricted 3D detector (its depthresolution is at most the thickness of the detector layer). If detectorsoffer 3D spatial resolution capability then interaction tracking(including multiple interactions) can be implemented internally within a3D detector layer as well as between detector layers.

Energy resolution can be used to measure the position-dependent energylosses due to the interactions within detectors which in turn canprovide an estimate of the energy of the initial incident x-ray or gammaray. This information can be used to determine whether the initialincident photon energy is within an allowed energy range as well as itsdirectionality.

Temporal resolution capability can be used to distinguish betweenindependent incident x-rays or gamma rays interactions (as well as theirsubsequent interactions) within the Compton camera. It will be shownthat very good temporal resolution can be beneficial if coincidencetiming is of interest between Compton camera systems (for example, whencoincidence PET imaging is implemented).

One implementation of a Compton camera using a dual-layer detectordesign wherein the first layer (front-end) was a small area, face-on, Sior Ge semiconductor pixelated detector offered 2D spatial resolution.The second layer (back-end) was a large area, face-on, scintillator(gamma camera) detector which also offered 2D spatial resolution (Singh,M., Medical Physics Vol. 10(4), pp. 421-427 (July/August 1983) andSingh, M., Doria D., Medical Physics Vol. 10(4), pp. 428-435(July/August 1983)). Both front-end and back-end detectors offeredappropriate levels of energy resolution for the photon energies employedand temporal resolution for the expected event interaction rates.

Since Compton scattered photons include a range of scatter angles thesensitivity of design is in part dependent on the separation distanceand area of the second layer with respect to the first layer ofdetectors. A second layer which employs a smaller 3D detector may, insome instances, be more-cost effective than a larger 2D detector whichsuffers from parallax errors and needs to be positioned further awayfrom the first layer.

Another implementation of the Compton camera, the (face-on) Comptontelescope camera, consisted of only a first layer detector. Thisfront-end detector was comprised of a stack (and thus could also beviewed as a multilayer detector) of 2D, face-on Si detectors whichfunction together as a 3D detector (Kroeger R, et al., IEEE Trans. Nucl.Sci., Vol. 49(4), pp. 1887-1892 (August 2002); Nelson, U.S. Pat. No.8,017,906). A different single layer (mono-layer) design implemented acylinder-like, 3D Ge detector defined by a positional readoutimplemented on the periphery and the hollow core of the detector.

A stack of 2D, face-on Ge detectors (or a thick 3D Ge detector with DOIcapability) can also be implemented although the Ge semiconductoroperates at a low temperature. The Compton telescope camera tracksmultiple Compton scatters by a photon in order to determine its originaldirection and energy.

Low-Z (atomic number) semiconductor materials such as Si and diamond(and sometimes moderate-Z Ge) are often preferred for the front-endCompton scatter detector for photons of relatively low energies (e.g.medical diagnostic x-ray energies, 140.5 keV gamma rays from Tc-99m usedin nuclear medicine) compared to the 511 keV gamma rays used in PETimaging.

The Compton scatter interaction cross section of the material dominatesits photoelectric cross section and the relative contribution to theangular reconstruction error due to the Doppler shift is reduced as Zdecreases and/or photon energy increases. As the photon energy increasessemiconductor materials with moderate-Z values (such as Ge, GaAs, CdTe,CZT, etc.) represent increasingly acceptable substitutes for low-Zsemiconductor materials such as Silicon.

The amount of energy deposited by relatively low energy photons(commonly used in diagnostic x-ray imaging or nuclear medicine) due to aCompton scatter interaction is typically small and thereforesemiconductors detectors are employed as front-end detectors because oftheir superior energy resolution compared to most scintillatordetectors. In the dual-layer Compton camera design lower-cost 2Dscintillator detectors may be employed in place of semiconductordetectors as back-end detectors if they offer suitable spatial, temporaland energy resolution and stopping power.

The semiconductor front-end detector may be replaced by a lowtemperature front-end detector or by a scintillator (or gas) front-enddetector although energy resolution may suffer. Any significantreduction in accuracy of the calculated incident photon directionalityby Compton reconstruction algorithms can be augmented or supplanted byadditional information such as coincidence between detectors (used incoincidence PET imaging).

Compton electron tracking in a gas detector can be implemented althoughthis is typically very time-consuming. Cherenkov radiation, despite therelatively weak optical signals, can be exploited for time-of-flight(TOF) measurements. (Cherenkov radiation can be detected when generatedin optically-transparent mediums including fluids such as liquids andgases, scintillators and non-scintillators such as transparent plastics,glasses, fibers, diamond films, etc. Thus, transparent dielectricmediums other than scintillators and gases can be also be employed asCompton scatter or photoelectron detectors within a Compton cameradetector system although energy resolution could suffer based on thedetection of Cherenkov radiation alone. Inexpensive dielectric materialsmay be acceptable for those applications in which radiation scatterwithin the object is of reduced importance and therefor lower detectorenergy resolution is acceptable. Variations of detector designsdescribed herein can include measuring only a Cherenkov signal or aCherenkov signal and a fluorescence signal or an electronic signal.)

Potential advantages of this dual-layer implementation of a multilayerdesign may include a less-expensive front-end detector and/or afront-end detector that offers a feature such as fast (greater than 1nanosecond) or very fast (less than 1 nanosecond) temporal resolution.Very fast temporal resolution is of interest for TOF PET. Benefits ofTOF PET include improved image resolution and lower patient radiationdose. Furthermore, the use of coincidence information can also simplifythe requirements of the back-end detector.

Compton electron tracking can also be implemented within a detectorlayer and between detector layers that employ at least one ofscintillator-photodetector detector, semiconductor, structured and lowtemperature detectors. Since electrons readily interact with matterelectron tracking is preferably implemented when detecting energeticphotons which are Compton scattered, typically generating more-energeticelectrons with a more-directional nature. (A similar concept applies toenergetic photoelectric interactions which typically generatemore-energetic photoelectrons with a more-directional nature. Thus, aCompton camera could utilize sufficiently energetic photoelectricinteractions for image reconstruction by tracking the highly directionalphotoelectrons.)

The tracking of Compton scattered electrons as well as Compton scatteredphotons can be simplified by enabling longer path lengths for thescattered particles, improving the estimates of scattering angles.Examples of relatively thin, edge-on detector configurations thatincorporate gaps between adjacent detectors (including partially- orcompletely-offset detectors) are shown in FIGS. 1, 3 and 5.

Face-on detector configurations with gaps between detector layers canalso be implemented. Compton camera image reconstruction can be improvedif both the Compton scattered photon and electron are tracked since thesolution can be limited to a fraction of a cone surface rather than thefull cone surface.

The flexibility of the Compton camera design can be understood byconsidering front-end (single layer) detector and front-end withback-end (dual-layer) detector implementations of multilayer, edge-ondetector Compton camera designs which can be used for low energy andhigh energy photon imaging. In one dual-layer implementation thefront-end detector is used to detect low energy x-rays or gammas and theback-end detector acts to detect higher energy gammas as an edge-onSPECT/gamma camera or PET camera (Nelson, U.S. Pat. No. 7,291,841).Front-end detectors and back-end detectors can be differentiated basedon functionality and/or position. The front-end and back-end detectorsshould have at least one different property such as position, size,geometry (planar, box, partial-box, ring, partial-ring, etc.),directionality (focused, unfocused), spatial resolution, temporalresolution, energy resolution, interaction probability (materialdensity, thickness, interaction coefficients), orientation (edge-on,face-on, tilt), noise characteristics, detector operation (integrator,PC, PCE), etc.

A multilayer detector can include one or more front-end detectors andback-end detectors. Detector properties within a detector layer can beuniform or non-uniform (continuous, discontinuous, mixing one or more ofdetector materials, detector operation capabilities, detectororientations, detector temporal characteristics, etc.). A special caseof a multilayer detector is a single detector layer that incorporatesone or more front-end detectors and back-end detectors. This can beimplemented in structured detectors (such as edge-on structuredsemiconductor detectors including structured 3D semiconductor detectorsand structured mold semiconductor detectors, structured scintillatordetectors including 3D edge-on or face-on stacked cross-coupledscintillator rod detectors, multilayer scintillator block detectors,scintillating fiber bundle detectors, straw array detectors, etc.).

For example, stacked cross-coupled scintillator rods can vary thescintillator rod properties (material, interaction probabilities,density, temporal characteristics, brightness, etc.) as a function ofdepth (as well as within a layer and even within individual rods).Front-end stacked cross-coupled layers might use, for example, ascintillator(s) preferred for lower energies encountered in SPECT or avery fast scintillator suitable for TOF PET) while back-end stackedcross-coupled layers might use a scintillator(s) preferred for moderateor fast or very fast PET. Furthermore, scintillator rod properties canbe varied within at least one of a rod, a layer, between cross-coupledlayers. Varying scintillator temporal characteristic as a function ofposition could be used to improve event localization based on bothoptical signal sharing and different temporal decay characteristics ofscintillator rods.

Other means of event localization, including signal wave form analysisbased on calibration of the detector volume using at least one of thedirect (event) signal, reflected signals, cross-coupled signals andwavelength shifted signals (including reflected signals) can also beemployed. For example, since the effective rod length is approximatelydouble the physical rod length when the paths (including scatter) of thedirect signal and the reflected signal are evaluated this informationcan be incorporated into a calibration procedure. Event localizationinformation can be used to correct for optical signal (includingCherenkov radiation) propagation time in TOF PET for cross-coupledscintillator rods (and crossed-fiber scintillator rods) involving atleast one of the direct event signal, a reflected event signal from therod end opposite the photodetector readout, a cross-coupled eventsignal, a reflected cross-coupled event signal, a reflectedcross-coupled signal, a wavelength shifted signal (including a reflectedwavelength shifted signal), indirect signals and reflected indirectsignals. In the case of a cross-coupled event signal commonly employedWLS materials (with nanosecond response) may be problematic unless thedelays can be accurately calibrated.

Very fast WLS materials (including quantum dots), if available, can bedeployed otherwise other techniques to direct the cross-coupled eventsignal to the photodetector readout should be implemented. Cross-coupledscintillator rods incorporate aspects of single-sided and two-sidedreadout scintillator rods. The use of event localization information toimprove TOF information can be employed with other detector geometriesdescribed herein (for example, a transparent layer coupled to one ormore layers of scintillator rods, a pixelated layer coupled to one ormore layers of scintillator rods, scintillator sheets with imposed rodand/or pixel and/or intermediate layer structures, scintillating fibers,scintillator rods with a single-sided or two-sided readout, etc.).

Consider a planar or ring multilayer detector geometry with two(discontinuous) detector layers in which adjacent 3D edge-on silicondetector modules with PCE capability in the front-end and back-enddetector layers are tilted with respect to one another to achieve afocused detector geometry with respect to diverging radiation from atleast one source, with the adjacent detector modules in the back-enddetector layer offset to fill gaps between the adjacent modules in thefront-end detector layer and define a substantially continuous detectorconfiguration. (Optionally, these two layers can be treated as a singledetector layer.)

Furthermore, consider a multilayer detector with three detector layers(treat the two focused 3D edge-on silicon with PCE capability detectorlayers as a single detector layer followed by a 2D face-on scintillatorwith integration capability followed by a 3D edge-on scintillator withPCE capability) employed as a PET camera and x-ray CT imaging system.The 3D edge-on silicon layer and 2D face-on scintillator layer bothfunction as the front-end detector for CT (experiencing different energyspectrums) and alternatively one or both layers could be employed in adedicated CT imaging system. The 3D edge-on silicon layer also functionsas a front-end detector for PET (detecting gammas or scattering gammas).The 3D edge-on scintillator layer acts as the back-end detector for PET(detecting non-scattered gamma rays and scattered gamma rays due to the3D edge-on silicon layer).

A focused, edge-on Compton camera design was described that can employone or multiple (of the same or different materials) detector layers aswell as implementing additional features such as the offset (complete orpartial) of adjacent (neighboring) detector modules within a layer.Completely offset detector modules can be used to create two or moredetector layers (offset layers) which when employed together canapproximate a continuous detector (and therefor can be referred to aseither a single layer or two layers (front-end and back-end layers) ofdetector modules). The offset layer feature of an edge-on Compton cameradesign can be implemented in PC, PCE and energy integration versions ofdiagnostic CT detector, including ring and cone beam CT as well astomosynthesis, PET, CT-PET, Compton-PET, Compton-PET-CT, gamma camera,etc. (as described in Nelson, U.S. Pat. No. 7,291,841; U.S. Pat. No.7,635,848; U.S. Pat. No. 8,017,906; U.S. Pat. No. 8,115,174; U.S. Pat.No. 8,115,175; and U.S. Pat. No. 8,183,533; Danielsson, U.S. Pat. No.8,183,535; and Bornefalk, U.S. Pat. No. 8,378,310). This complete orpartial offset feature can be used for not only edge-on detectorimplementations but also face-on detector implementations for ring andcone beam CT (for example, a planar or cylindrical arrangement of lineararrays of face-on detectors, each oriented parallel to the axialdirection of the scanner) as well as tomosynthesis. It should beunderstood that the modifications and improvements described herein forring CT detector implementations are also applicable for cone beam CTand tomosynthesis detector implementations.

Implementations of the Compton camera design are described herein thatare suitable for use as Compton-PET imaging systems and CT-Compton-PETimaging systems. In addition, the positioning of nuclear medicinecollimator hardware such as focused, parallel or pin hole collimatorsbetween the object being imaged and the Compton camera permits thesystem of collimator and Compton camera to provide nuclear medicineimaging capabilities (the imaging capabilities of a SPECT/Gamma camera)for those applications in which the Compton camera does not offeradequate imaging properties.

Limited implementations of the Compton camera designs described hereininclude versions that function only as CT or PET (and SPECT) detectordesigns. The Compton camera imaging systems described herein will finduse in diagnostic medical x-ray CT, nuclear medicine and PET imaging,x-ray micro-CT imaging, dental CT, medium and small animal imaging,radiation therapy imaging, industrial imaging, HLS and military imaging,and scientific imaging.

Compton-PET Detector Systems

One implementation of the Compton camera is referred to as theCompton-PET detector system (Nelson, U.S. Pat. No. 7,291,841). TheCompton-PET detector system design allows flexibility in the choice ofdetector materials as well as detector geometries. This flexibility isconstrained by the intended imaging applications (such as PET only,nuclear medicine and PET, x-ray and PET).

Face-on, edge-on, and combinations of face-on and edge-on detectors canbe employed. One, two or more than two layers of detectors can beemployed. Detector modules that comprise a detector layer can optionallybe partially-offset or completely-offset from their adjacent neighborswithin a layer.

PET image acquisition formats based on planar and focused structure(such as ring and or partial ring) geometries are implemented.Compton-PET detector systems are based on block, 1D, 2D or 3D edge-on,face-on, or mixtures of edge-on and face-on detectors (including edge-ondetectors with SAR capability and face-on detectors with DOI capability)(Nelson, U.S. Pat. No. 4,560,882; U.S. Pat. No. 4,937,453; U.S. Pat. No.5,258,145; U.S. Pat. No. 6,583,420; U.S. Pat. No. 7,291,841; U.S. Pat.No. 7,635,848; U.S. Pat. No. 8,017,906; U.S. Pat. No. 8,115,174; U.S.Pat. No. 8,115,175; and U.S. Pat. No. 8,183,533). The non-coincidenceand coincidence Compton-PET detector systems described herein includefocused and unfocused planar detector formats and focused structure(such as ring and partial ring as well as focused ring and focusedpartial ring) detector formats.

A non-coincidence Compton-PET (one-sided PET) detector system isimplemented by extending Compton camera designs that have been developedfor nuclear medicine imaging devices such as hand held probes orSPECT/Gamma cameras so that the detector system can operate with thelower gamma ray energies used in nuclear medicine as well as the higherenergy range of PET with good detection efficiency. A highly flexibleimplementation of a Compton camera design is a dual-layer, 3D Comptoncamera. A specific implementation, a non-coincidence Compton-PETdetector system, employs a (preferably, but not exclusively) Comptonscattering front-end detector and a (preferably, but not exclusively)high-stopping power back-end detector in which both front-end andback-end detectors offer suitable 3D spatial resolution, energyresolution and temporal resolution (Nelson, U.S. Pat. No. 8,017,906).

Both the front-end and back-end 3D detectors provide adequate temporalresolution for an expected event rate, such that accurate event trackingcan be enabled both within the front-end and back-end detectors andbetween the front-end and back-end detectors. Both the front-end andback-end 3D detectors can record Compton scatter and photoelectricinteractions.

In some instances Raleigh scattering interactions (angle change withinsignificant energy loss) can be identified based on trackinginformation. The front-end and back-end detectors, either separately ortogether, can operate as two layer Compton cameras and Compton telescopecameras (Nelson, U.S. Pat. No. 8,017,906).

In one scenario the 3D front-end detector can function as a single (ormultiple) Compton scatter device and the 3D back-end detector can beused to measure the energy and interaction location of the Comptonscattered photon. The front-end and back-end detectors have 3D spatialresolution. Front-end and back-end 3D detectors can also Compton-scattera photon (measuring position and energy deposited) and detect the(single or multiple) Compton-scattered photon (measuring its energy andinteraction location). Therefore this two layer Compton camera with 3Ddetector layers incorporates the capabilities of three two layer Comptoncameras (in which one layer Compton-scatters the photon and the otherlayer detects (stops) the Compton-scattered photon).

Compton telescope camera designs exploit multiple Compton scattering forreconstruction. The Compton telescope camera capability can beimplemented in the 3D front-end detector, in the 3D back-end detectorand between the 3D front-end and back-end detectors (providing thecapabilities of three (multilayer, face-on 2D array detectors) Comptontelescope cameras).

Appropriate two layer Compton camera and Compton telescope camerareconstruction algorithms are used to form an image. When this Comptoncamera is used to image single annihilation gamma rays created during aPET scan it is referred to as a one-sided PET detector system or anon-coincidence Compton-PET detector system. (This dual-layer, 3DCompton camera design is clearly not limited to PET imaging alone andtherefore may be adapted for use in imaging applications at other photonenergies. Furthermore more than two layers of 3D detectors can beemployed and non-3D layers of detectors can be mixed with 3D layers ofdetectors, thereby introducing additional flexibility in the types ofimaging applications for which this Compton cameras design is suitable.)

This one-sided PET detector can be implemented in a focused or unfocusedplanar detector geometry or a focused structure detector geometry suchas a ring or partial ring (as well as focused ring and focused partialring detector geometry). This avoids the expense of employing acoincidence PET detector system based on opposing (or nearly-opposing)sets of PET detectors.

EXAMPLES

FIG. 1 illustrates a dual-layer Compton-PET detector imaging system 1000that incorporates 3D, edge-on detector arrays 510 and 520 (a first layerof detectors and a second layer of detectors, respectively). Theindividual, 2D edge-on detector modules 102 use crossed strip radiationdetectors 115. Alternatives include 2D pixelated arrays (or 3D pixelatedarrays if SAR capability is enabled) in an edge-on geometry.

Incident radiation photons 107 from gamma ray radiation source, withenergy less than the pair production threshold, can undergo Rayleighscattering, Compton scattering or photoelectric interactions. Comptonscattered gamma ray photons 108 can be detected by the edge-on radiationdetector within the module 102 responsible for the initial scattering,by other edge-on detector modules within the front-end detector layer510 (detector layer 1) or by detector modules within the back-enddetector layer 520 (detector layer 2).

Each module 102 also includes a base 106 and a communications link 103.The base 106 preferably contains detector electronics including signalconditioners and readout ASICs, power management components, temperaturecontrol components, and a data or information channel for communicatingwith the computer system. The communications link 103 can be used toprovide power to the module 102 and connects the base 106 to a computersystem.

The communication link 103 preferably is used to off-load the digitizeddetector radiation data to a computer system for analysis and imagereconstruction. The computer system, which can include general purpose,dedicated, and embedded computers, provides monitor and control servicesto modules 102, to the detector layers 510 and 520 and to the entireCompton-PET detector imaging system 1000.

The computer system evaluates module parameters, detector layerparameters, and the detected radiation image data. The detected data isprocessed and can be displayed and stored if desired. Additionalrelevant module information, such as temperature, amplifier settings,detector voltages, position, orientation, and motion information, can betransmitted to this computer system over the communication link 103. Thecomputer system transmits instructions that update the detector modules102 and detector layers 510 and 520. This establishes a dynamicinformation feedback loop that is useful for adaptive imaging (Nelson,U.S. Pat. No. 7,291,841).

Note that the electronic functionality of the detector base 106 can beimplemented along the side of a detector module or attached to thesurface of the detector module (integrated electronics). Another optionwhen the detector substrate is a semiconductor such as Si is to etch anindentation along the bottom of (opposite the radiation entrancesurface) and mount the readout ASICs and radiation shielding in theindentation and directly to the substrate along the bottom edge.

If the length of the edge-on detector is greater than its height thenthis configuration allows the readout ASICs to be closer to a set ofdetector pixels than for the case wherein the readout ASICs are mountedalong the side in order to avoid the direct x-ray beam. Preferably thecombined thickness of the etched substrate and mounted readout ASIC withshielding would be less than or equal to the thickness of the substrate(avoiding problems if the readout ASIC and any shielding stick out fromthe substrate and possibly interfering with the x-ray beam seen byoffset detectors).

FIG. 2 illustrates a perspective of readout ASICs 200 with radiationshielding 204 mounted in an etched Si substrate 208 (or another suitablesemiconductor substrate), with a pixel size 215 that varies with heightwhich is positioned edge-on to incident radiation photons 109. Othertechniques of delivering power to the detector modules as well aswireless communication can be employed in place of communication link103 (FIG. 1). It should be understood that readout ASICs can be mountedalong the side and the bottom edge.

Two or more non-coincidence Compton-PET detector systems (an enhancednon-coincidence Compton-PET detector system) can be employed for a PETimaging application. Furthermore, with the addition of coincidencecircuitry, pairs of non-coincidence Compton-PET detector systems(preferably facing each other and positioned on opposite sides of anobject) can function as a coincidence Compton-PET detector system.

The cost of a two layer non-coincidence Compton-PET (one-sided PET)detector system can be reduced if either one or both of the 3D front-endand back-end detectors can be replaced by a suitable 2D detector withacceptable energy and temporal resolution and stopping or scatteringpower. The caveat is that photon detection efficiency and reconstructionimage quality may suffer as a result. A compromise in terms of cost isto leave the front-end detector with 3D spatial resolution (andtherefore retaining the previously herein-listed capabilities: tofunction as a Compton scatterer, a two layer Compton camera, a Comptontelescope camera), and employ a back-end detector with 2D spatialresolution. The back-end detector would offer acceptable stopping power,energy resolution and temporal resolution for the expected gamma rayevent rate and gamma ray energies.

For a planar detector geometry the front-end and back-end detectors canconsist of single-layer face-on detector plane modules, a multilayer(stack) of face-on detector plane modules, a single-layer of edge-ondetector modules, a stack of edge-on detector modules or a combinationof face-on and edge-on detector modules. Face-on detector modules caninclude DOI capability whereas edge-on detector modules can include SARcapability.

One implementation of a focused planar detector geometry (suitable forcone beam CT, tomosynthesis, etc.) employs a front-end detector thatconsists of either a single layer (offset or non-offset) or multiplelayers (offset or non-offset) of tilted edge-on (and/or face-on)detector modules. A degree of physical focusing (promotingdirectionality) is achieved by tilting the detector modules (detectormodules with fixed or adjustable tilt angles can be implementeddepending on the imaging requirements). As an alternative to a parallelpixel structure a focused pixel structure can be implemented along thelengths of the edge-on tilted (or parallel) detector modules to accountfor x-ray beam divergence (which can also be implemented in a ring orpartial ring CT detector geometry).

Furthermore, an additional degree of physical focusing can be achievedby positioning detector modules (using parallel and/or focused pixelstructures) in a curved geometry and thereby approximating arc-shapeddetector lines (suitable for a focused, near-planar detector geometry aswell as ring or partial ring CT detector geometries). Each of the offsetor non-offset edge-on detector module comprising the first layer oftilted edge-on detector modules can have at least a second (offset ornon-offset) edge-on or face-on detector module (comprising the at leastsecond layer of detector modules), tilted or not tilted, positionedbeneath it. For example, the first layer can implement offset tiltededge-on silicon detector modules with each offset silicon detectormodule followed by one or more semiconductor or scintillator face-on oredge-on detector modules comprising one or more additional layers(typically employing moderate-to-high Z detector materials).

FIG. 3 illustrates a perspective of a focused planar detector system1000 in which detector modules 102 are tilted so as to focus ondiverging radiation 109 from a radiation source. In addition the pixelstructure 115 within the individual detector modules 102 is angled so asto focus on the same radiation source.

The tilting of the detector modules may create unacceptable gaps betweenneighboring detector modules within the detector layer 510. These gapsare shown to be effectively filled by the complete offset of every otherdetector module comprising the offset detector layer 510.

One implementation of a focused structure detector geometry such as aring (or partial ring) employs a front-end detector comprised of asingle layer (non-offset) or single layer with an offset layer (whichcan be treated in this application as a single layer) of tilted edge-ondetector modules. As in the case of planar detectors, a focused pixelstructure can be implemented along the lengths of the edge-on tilteddetector modules (creating focused ring and focused partial ringdetector geometries).

Suitable detector configurations and materials have been described forCompton, PET, nuclear medicine and x-ray imaging (Nelson, U.S. Pat. No.6,583,420; U.S. Pat. No. 7,291,841; U.S. Pat. No. 7,635,848; U.S. Pat.No. 8,017,906; U.S. Pat. No. 8,115,174; U.S. Pat. No. 8,115,175; U.S.Pat. No. 8,183,533; U.S. patent application Ser. No. 13/199,612 (U.S.Publication No. 2012/0181437); and U.S. Pat. No. 9,347,893). Examples ofsuitable detector configurations include a single or multilayer face-ondetector, a single or multilayer edge-on detector and a multilayerdetector comprised of face-on and edge-on detectors.

Edge-on detectors may incorporate SAR capability and face-on detectorsmay incorporate DOI capability. Examples of suitable detector materialsand formats include semiconductor detectors, structured detectors suchas single and double sided structured 3D silicon as well as otherstructured 3D semiconductor materials (Diamond, Ge, Se, GaAs, CdTe, CZT,etc.), structured quantum dots, structured scintillators, andscintillators. Structured mold quantum dot detectors (also referred toas structured quantum dot detectors) offer flexibility since a varietyof cell shapes (including trenches) can be implemented. Furthermore, theselection of (and density of) quantum dot (nanoparticle) materials canbe varied as a function of position within the substrate in order toenhance a type of interaction such as Compton scattering or thephotoelectric effect. Silicon is frequently used as a mold material inthe form of porous silicon or micromachined silicon for semiconductorquantum dots. Silicon and other mold materials can be used withscintillator quantum dots as well as scintillator materials.

Structured mold semiconductor detectors implement (but are not limitedto) either semiconductor quantum dots (nanoparticles) or amorphoussemiconductors or polycrystalline semiconductors (semiconductormaterials). The flexibility of the structured mold architecture enablesincorporating not only two or more semiconductor materials within astructured mold but also implementations such as one or moresemiconductor materials with one or more scintillator materials and/orgases, one or more scintillator materials with one or more gases, etc.within a structured mold. For example, an edge-on dual-layer detectorwith a semiconductor detector first (front-end) layer and a scintillatordetector second (back-end) layer can be manufactured as a single,edge-on structured mold detector with semiconductor and scintillatorcomponents.

The first layer within the structured mold could implement one or moresemiconductor quantum dot (nanoparticle), amorphous semiconductor and/orpolycrystalline semiconductor materials in appropriate geometries (inthis implementation the first layer is comprised of one or more layers)for the incident radiation field. The second layer could implement oneor more organic and inorganic scintillator materials including, but notlimited to, scintillator quantum dot (nanoparticle), polycrystallinescintillator, nanophosphor scintillator, liquid scintillator, gasscintillator, etc. materials in appropriate geometries (in thisimplementation the second layer is comprised of one or more layers) forthe incident radiation field.

Partial lists of suitable organic and inorganic scintillators andsemiconductors are provided, e.g., in Knoll G., Radiation Detection andMeasurement, 4th edition, Wiley (2010) which is incorporated byreference for these teachings as described herein. Suitable materialsinclude, but are not limited to, organic crystal scintillators,inorganic crystal scintillators, plastic (polymer) scintillators and(plastic and non-plastic) scintillating fibers and fiber bundles(strips) (scintillating fiber bundles (strips) represent oneimplementation of a structured detector), gel scintillators, liquidscintillators, deuterated liquid scintillators, and loaded liquidscintillators (loaded, e.g., with B, Gd or Sn). Suitable gasscintillators include, but are not limited to, xenon, krypton, argon,helium, and nitrogen. Glass scintillators may also be used (e.g.,silicate glass containing lithium activated with cerium).

Additional detector options include structured, gas-filled strawdetectors with appropriate low-Z or moderate-Z material annuli whichprovide suitable energy, spatial and temporal resolution and stopping orscattering power (Nelson, U.S. Pat. No. 8,017,906), liquefied gas baseddetectors (such as Xenon), semiconductor-based or gas-based Medipixdetectors and low temperature (such as GE and superconductor) detectors.Multiple Compton-PET (one-sided PET) views of a volume of an object tobe imaged can be acquired as a result of detector system rotation aboutthe object to be imaged.

An alternative imaging format is to rotate the object and keep thedetector system stationary. Additional object volumes can be imaged, ifneeded, by translating (typically) the object through the scannersystem.

It should be noted that if the Compton camera image quality isn'tsuitable for the nuclear medicine imaging applications of interest thena collimator can be inserted in front of the detector so that the systemof collimator and detector can function as a SPECT/gamma camera. Sincethe collimator imposes a degree of directionality then the SPECT/gammacamera implementation of a Compton camera can utilize both Comptonscatter interactions (and tracking capabilities) as well as directphoto-electric interactions (which have a much higher probability ofoccurring at lower energies such as 140.5 keV versus 511 keV in low-Zand high-Z detectors). The direct photo-electric interactions would notbe used in conventional (no electron tracking) Compton camera imaging.Furthermore, a miniature version of the Compton-PET detector system canbe implemented as a Compton-PET hand-held detector probe. The additionof a nuclear medicine collimator permits the Compton-PET detector probeto function as a gamma camera hand-held detector probe. Versions ofprobes can be operated in non-coincidence or coincidence mode withnon-coincidence Compton-PET detector systems (as well as coincidenceCompton-PET detector system) to offer enhanced resolution.

Coincidence Compton-PET detector systems extend the implementations of anon-coincidence Compton-PET detector system by including a secondCompton-PET detector system and coincidence circuitry between the pairof Compton-PET detector systems, for example, employing a pair of planaror partial ring Compton-PET detector systems with coincidence circuitry.

FIG. 4 illustrates a coincidence Compton-PET detector system which iscomprised of a pair of planar Compton-PET detector systems 1000 withcommunications links 103 operated in coincidence for imaging an object111 (for example, the heart). Each planar Compton-PET detector system1000 is positioned by an electronically controlled actuator arm 130.

For the case of a partial ring Compton-PET detector system, if asufficient number of pairs of partial ring Compton-PET detector systemsand coincidence circuitry (linking all detectors) are employed, then acomplete ring coincidence Compton-PET detector system can beimplemented. The complete ring geometry can be achieved with a singlepair of partial ring Compton-PET detector systems if each partial ringcovers an angular aperture of 180 degrees.

If the Compton scatter capability of a front-end detector is not needed(for example, if only one complete Compton camera is needed for non-PETimage applications), then there is the option of employing only aPET-compatible detector for the second detector system. Additional pairsof Compton-PET and/or PET-compatible (or combinations of both) detectorswith appropriate coincidence circuitry can be combined to form anenhanced coincidence Compton-PET detector system. (Note that a dummy ornon-functional equivalent of the front-end detector can be used to makea stand-alone PET-compatible detector unit “see” a comparable radiationfield to what the back-end detector experiences in a coincidenceCompton-PET system without the cost of an active front-end detector).

The description of a flexible non-coincidence Compton-PET detectorsystem applies to the Compton-PET detector systems used in a coincidenceCompton-PET detector system. Consider the case in which at least one ofthe two detector system is a Compton-PET detector system. The front-endand back-end detectors offer suitable 3D spatial resolution, energyresolution and temporal resolution and stopping or scattering power.Both the front-end and back-end detectors provide adequate temporalresolution for an expected event rate such that accurate event trackingcan be enabled both within the front-end and back-end detectors andbetween the front-end and back-end detectors, since Compton scatter andphotoelectric interactions can be recorded in both front-end andback-end detectors.

As described for non-coincidence Compton-PET detector systems thiscombination of front-end and back-end detectors incorporates thecapability of three two-layer Compton cameras and three Comptontelescope cameras. The addition of coincidence detection capabilityintroduces additional flexibility in that events involving a singlephotoelectric interaction (in which no Compton scattering occurs) in thefront-end or back-end detector can be used for coincidence detection aswell as events involving one or more Compton scatter interactions.

In a conventional Compton camera design a photoelectric interaction inthe front-end detector layer is not useful. A fast or very fast detector(including, but not limited to, silicon, GaAs, various structured 3Ddetectors such as 3D silicon, structured mold, etc.) can provide timinginformation for coincidence PET using either photoelectric or Comptonscatter interactions. The capability to use photoelectric events withoutCompton scattering leads to an alternative detector system design inwhich the front-end detector layer employs a moderate-to-high Z detectormaterial with tracking capability. In this implementation bothphotoelectric events and Compton scattering can be exploited but now thephotoelectric interaction relative probability is more significantcompared to a material such as silicon. Tracking capability for Comptonscattered photons (as well as characteristic x-rays) can be used forestimating the deposited energy for each detected event even if Comptonscatter reconstruction is not employed. If the front-end detector issufficiently fast then the TOF PET imaging can be implemented. Forexample, an edge-on, structured mold detector implementing at least oneof high-Z (semiconductor) quantum dots, amorphous semiconductors, andpolycrystalline semiconductors can replace an edge-on silicon orstructured 3D silicon detector. If the edge-on, structured mold detectoroffers significant attenuation it can be used in place of both thefront-end and back-end detectors. Furthermore, this alternative detectorsystem design can be readily extended for use with CT-Compton-PET,CT-PET, PET Compton-PET and CT detector systems.

Since very fast coincidence timing (TOF) can be used to improvereconstruction accuracy and reduce patient dose and/or image acquisitiontime there can be a benefit from having one or both of the front-end andback-end detectors capable of very fast timing resolution. If bothfront-end and back-end detectors are involved in the detection processthen coincidence timing can be based on using at least one of thefront-end and back-end interaction timings. Timing resolutioncorrections are made for the response of one or both detectors(depending on whether one or both of the front-end and back-enddetectors are involved in detection) and gamma ray travel times betweeninteraction locations within one or both detectors and between detectors(Nelson, U.S. Pat. No. 8,017,906).

Commercial TOF PET systems are capable of very fast temporal resolution(on the order of or less than one nanosecond). Very fast temporalresponse capabilities can influence the choice of detector materials forfront-end and back-end detectors. If the front-end detector has areasonable probability per photon of a Compton scatter interaction thenone option is to select a front-end detector material with a very fasttemporal response and select a (possibly much less expensive) back-enddetector material with a much slower temporal response.

If a gamma ray undergoes a Compton scatter interaction in at least oneof the front-end and back-end detectors as well as additionalinteractions such that the energy of the incident particle can beestimated, then photon directionality based on the appropriate Comptoncamera reconstruction algorithm (for the Compton camera designsdescribed for non-coincidence Compton-PET detector system) can becompared with photon directionality based on coincidence (line-of-sight)between the Compton-PET detector systems operating in coincidence. TheCompton-based directionality can be used to estimate the degree ofvalidity of the coincidence (line-of-sight) assumption, includingacollinearity. This capability can be used to help reject some of thephotons that undergo Raleigh and/or Compton scattering within the objectand its surroundings as well as Rayleigh scattering or difficult todetect Compton scattering within the detectors.

In addition, a (combined) non-coincidence Compton-PET (one-sided PET)reconstructed image can be compared to (or combined with) a coincidencePET reconstruction image. (Nelson, U.S. Pat. No. 8,017,906). Unpaireddetected events (in which coincidence fails since only one of the twoannihilation photons is detected and is considered legitimate) by aCompton camera can still contribute to the Compton scatterreconstruction image.

As described for the case of non-coincidence Compton-PET (one-sided PET)detector systems, system cost (in some cases) may be reduced if theback-end detector 3D spatial resolution capability is lowered to 2Dcapability while maintaining adequate energy and temporal resolution.The 2D spatial resolution of the back-end detector implies that itoffers limited performance as a stand-alone PET detector for gamma raysthat aren't Compton scattered by the front-end detector.

The back-end detector should provide good stopping power. The Comptonscattering front-end detector offers suitable 3D spatial, temporal andenergy resolution and scattering interaction probability. Single andmultiple Compton scattering (as well as photoelectric) interactions canoccur in the front-end detector, allowing the front-end detector tofunction as a Compton camera, as a PET camera, as the first layer of amultilayer Compton camera and as the first layer in a multilayer PETcamera in which it records the initial interaction location, energydeposition and event timing information. (Note that if the multilayerCompton camera capability is sacrificed then the 2D spatial resolutioncapability of the back-end detector can be reduced to 1D or even blockdetector spatial resolution, further reducing costs. The back-enddetector primarily provides stopping power along with appropriate energyand temporal resolution. The front-end detector should offer anacceptable probability of undergoing at least one Compton scatterinteraction so that an initial location of interaction, timing andenergy deposition can be established. If TOF PET imaging is desired thenthe front-end detector can offer very fast temporal resolution. Thefront-end detector, due to its 3D spatial resolution capability, canstill track multiple scatter interactions as well as photoelectricevents. The front-end detector retains the capabilities of a Comptoncamera and a PET detector. Event tracking between the front-end andback-end detectors is employed.)

The back-end detector can also offer fast or very fast temporalresolution. The front-end detector can maintain fast or very fasttemporal resolution capability but an alternative is to implement afront-end detector (with suitable temporal resolution) primarily forestablishing spatial resolution via photoelectric and Comptoninteractions while relying on the fast or very fast back-end detector(with suitable spatial resolution) to establish coincidence timing forthe front-end detector Compton scattered photons. In someimplementations cost savings may be realized and the choice of front-endand back-end detectors may be expanded by moving some temporalresolution capabilities from the front-end detector to the back-enddetector. For example, the 3D front-end detector could use low, moderateor high-Z semiconductor materials with less emphasis on temporalresolution and more emphasis on spatial resolution while the back-enddetector could offer reduced spatial resolution while emphasizingtemporal resolution (employing fast scintillators such as LSO, LYSO,BaFl₂, LaBr₂, etc. coupled to PMTs, microchannel plates, SiPMs, siliconnanowires, etc.). Implementations of these dual-layer detectorconfiguration are suitable for use in CT-PET imaging systems wherein thefront-end detector is used for CT and PET and the back-end detector isused for CT and PET or PET. Furthermore, the front-end detector andback-end detector layers can always be implemented as independentdetector layers leading to cost savings. For example, the CT front-enddetector can be a simple face-on scintillator (energy integrator), adual-energy edge-on scintillator, a PCE edge-on semiconductor detector,etc. but the needed for capabilities such as tracking electronicsbetween layers, etc. are removed.

Multiple Compton-PET or PET views of an object volume to be imaged canbe acquired as a result of detector rotation about the object. Thealternative imaging format is to rotate the object and keep the detectorsystem stationary. If the Compton camera image quality isn't suitablefor the nuclear medicine imaging applications of interest then acollimator can be inserted in front of the detector so that the systemof collimator and detector can function as a SPECT/gamma camera(collimators can also be used with PET cameras).

For the coincident and non-coincident Compton-PET configurationsdescribed there are many options for detector materials based on costand performance requirements. Assuming that acceptable-to-good energyresolution is desirable, then block, 1D, 2D and 3D back-end detectorsand 2D and 3D front-end detectors can use semiconductors,polycrystalline and amorphous semiconductors, structured 3Dsemiconductors, structured mold semiconductor quantum dots(nanoparticles) as well as amorphous semiconductors and polycrystallinesemiconductors, moderate-to-bright nanophosphors (scintillator quantumdots), organic and inorganic scintillators, gas and liquid detectors,and amplified detectors. Furthermore these detectors can incorporateedge-on SAR or face-on DOI (positional encoding) capabilities.

Semiconductor and gas detectors typically offer a Fano factor noticeablyless than 1.0. If stopping power is important, then sufficient detectormaterial can be present in order to provide good to excellentattenuation. Detector response time (for example, scintillator decaytime) properties should be suitable for at least event tracking atexpected event rates. Very fast detectors would permit the use of TOFinformation to be utilized in PET reconstruction algorithms.

Possible scintillators with at least one of these properties include,but are not limited to: BaFl₂, LaBr₃ (including co-doped), Tl₂LiLaBr₆,LaCl₂, LSO, LYSO, GSO, GdI₃, LuI₃, SrI₂, BaHfO₃, SrHfO₃, PbWO₄, LuAP,CsI:Tl,Sm, NaI:Tl, BGO, CsI:Tl, Lu₂O₃:Eu, ZnO-based fast scintillatorsas well as glass, plastic and fiber scintillators, liquid scintillators,gas scintillators, quantum dot scintillators, ceramic scintillators,polycrystalline scintillators and various fast-to-very fast organicscintillators. Possible semiconductor detectors (and variants thereof)with at least one of these properties include, but are not limited to:diamond, Si, SiC, Se, Ge, GaAs, GaAs:Cr, CdTe, CZT, HgI₂, PbO, PbI₂,TlBr (as well as low noise implementations such as silicon driftdetectors or those with gain such as Si-APDs or SiPMs or siliconnanowires or iDADs, Se-APDs, GaAsPMs and DiamondPMs) detectors; onedimensional structures such as rods and wires, structured single anddouble sided 3D Si and other semiconductor material detectors andstructured mold semiconductor quantum dot (nanoparticle), amorphoussemiconductor, and polycrystalline semiconductor detectors.

A number of these semiconductor detectors can be configured as fast orvery fast photodetectors and so they can be coupled with fast or veryfast scintillators such as quantum dot, organic, or inorganicscintillators. Suitable detector packages (a detector material coupledto a readout ASIC) include Medipix-based detectors. Additionalstructured detectors with gain include, but are not limited to,gas-filled straw detectors (Nelson, U.S. Pat. No. 8,017,906).

In addition, the choice of detector material can be influenced by thedetector format. For example, a 10 mm thick (or greater) CdTe or CZTface-on detector (used primarily for photo-detection) for PET imagingmay offer limited temporal resolution, whereas a 1 mm thick (or less)CdTe or CZT edge-on detector (used for photo-detection and/or Comptonscattering) may qualify as a fast detector (even if SAR or DOIcorrections are not implemented). From a similar perspective a 1 mm or0.5 mm (or less) thick Si or Ge edge-on detector (used for Comptonscattering or Compton scattering and photo-detection) can be employed asa fast or very fast detector.

If SAR or DOI capabilities are implemented to estimate the interactionlocation of an event, then timing corrections can be made based on thepropagation times of electrons or holes to the anode and cathode,respectively (Nelson, U.S. Pat. No. 7,635,848; U.S. Pat. No. 8,017,906).An edge-on or face-on structured 3D semiconductor or structured moldsemiconductor quantum dot detector can be employed as a fast or veryfast detector since charge propagation distances can often be less than40-100 microns.

The flexibility of this Compton-PET design also allows alternativechoices for the front-end detector and back-end detector based onfactors such as lower cost and non-redundancy of features (if possible)as well as spatial resolution, energy resolution, temporal resolutionand the likelihood of Compton scatter and photoelectric interactions.For example, a Compton-scatter front-end scintillator detector could beemployed based on suitable (excellent) timing resolution and/or spatialresolution despite reduced (or minimal) energy resolution compared to asemiconductor detector. Compton reconstruction techniques, in someinstances, can be employed to estimate the photon energy loss due to thefront-end detector Compton scatter event.

Suitable front-end detector candidates with at least one of theseproperties include, but are not limited to: low-Z or moderate-Z, fastand very fast organic or inorganic scintillators (or scintillatingfibers) with a suitable high-speed, sensitive optical readout detectors(such as photodiodes, APDs, semiconductor photomultipliers such asSiPMs, silicon nanowires and GaAsPMs, electron multiplier CCDs,microchannel plates, etc.), semiconductor-based, scintillator-based orgas-based Medipix detectors, and structured, gas-filled straw, RPC, etc.detectors with appropriate low-Z or moderate-Z material which functionas a source of Compton electrons. In additional examples, straw andplastic (or non-plastic) scintillating fiber detector formats, whenimplemented in an edge-on geometry used for PET imaging, could implementSAR (permitting position estimates as well as energy and timingcorrections). Furthermore, plastic scintillating fibers (as well asnon-plastic scintillating fibers) can be coated with thin films ofmoderate-to-high-Z materials to enhance their photoelectric crosssection (permitting the properties of the front-end detector layer to betuned in terms of the Compton scatter and photoelectric interactionprobabilities).

Previously, structured straw (gas-filled) detectors incorporated onlyhigh-Z annuli in order to enhance the photoelectric effect (Nelson, U.S.Pat. No. 8,017,906). The same design technique can be used with low-Zand moderate-Z annuli in order to enhance the Compton scatter effect.Furthermore, combinations of low/moderate-Z annuli straw detectorsfollowed by high-Z annuli straw detectors (or other high-Z detectors)can be employed. Gas-filled detectors including, but not limited to,structured straw detectors and RPC detectors can be manufactured insizes appropriate for conventional PET or extended axial field of viewwhole body PET (potentially lowering detector cost).

Furthermore, the relatively high photon energies encountered in PETimaging favor forward scattering of Compton electrons. For a single RPCdetector the incident annihilation photons initially encounter a thinsheet of glass, plastic, bakelite, etc. in the upper surface, with arelatively small probability of undergoing Compton scattering and a muchsmaller probability of undergoing a photoelectric interaction. The uppersurface is the dominant source of ionizing electrons events.

An alternative to stacking single RPC detectors to form a multilayer RPCdetector is to construct a multilayer RPC detector in which the lowersurface of one RPC detector is utilized as the upper surface of the nextRPC detector in the stack (this has limits in terms of voltage divisionbetween plates). This can be a problem for a multilayer straw detectorwhich emphasizes Compton scatter events for PET imaging (whether it isoriented face-on or edge-on or at an intermediate angle with respect tothe incoming annihilation photons). For example, for a face-on geometrythe potential ionizing electron events generated within the bottomsurface wall of a straw must then penetrate the upper surface wall of astraw in the next layer in order to be detected. (Note that the strawdetectors walls can be manufactured from a variety of low-Z materials(glass, plastic, carbon, Al, etc.).

Detectors should offer an acceptable probability of experiencing atleast one Compton scatter interaction so that an initial location ofinteraction can be established. Event tracking within and between thefront-end and back-end detectors can be employed. If the front-enddetector offers excellent temporal resolution then TOF information canbe used to improve the reconstructed image along with a reduction inpatient dose and/or image acquisition time. If a front-end detectorlacks good energy resolution it still can be effective if the front-endand back-end detectors offer good spatial resolution and the back-enddetector offers good energy resolution.

Coincidence (line-of-sight or line-of-reaction) directionality can beexploited along with the scattered photon angle in order to estimate theincident gamma ray energy for cases of simple Compton scatter. Once theproperties of the front-end or back-end detector have been defined, thenthe properties of the other detector can be selected on the basis ofwhich properties need to be accentuated or can be allowed to diminish(such as stopping power, energy resolution, spatial resolution andtemporal resolution).

The back-end detector may primarily offer stopping power and energyresolution if the front-end detector offers 3D spatial resolution andenergy resolution. Then a cost-based decision can be made as to whetherthe front-end or back-end detector (or both) should provide acceptable,fast or very fast temporal resolution.

Thus a single detector implementation does not have to embody all of thecoveted PET detector properties (high stopping power, 3D spatialresolution, fast or very fast temporal resolution). For example, thecoincidence Compton-PET detector system can implement features such asTOF imaging with a range of detector options that is much greater thanwith commercial (conventional) TOF PET systems. Nonexclusive lists ofsuitable scintillator and semiconductor materials are provided herein.Partial lists of suitable organic and inorganic scintillators andsemiconductor materials including some of their properties are providedin Knoll G., Radiation Detection and Measurement, 4th edition, Wiley(2010), p. 230 (table 8.2), p. 238 (Table 8.3) and p. 492 (Table 13.3),respectively, each of which is incorporated by reference herein, in theentirety and for all purposes.

The flexibility of using front-end and back-end detectors for PET whichcan offer different spatial, temporal and energy resolution for PETresults in different PET images based on which detectors interact withthe pair of gamma rays from an annihilation event. For example, aCompton-PET front-end detector could Compton scatter one gamma of a pairwhich is then detected by the back-end detector. Another Compton-PETfront-end detector might fail to scatter the other gamma of the pairwhich is detected by the back-end detector. Coincidence can beestablished but the timing or spatial resolution (or both) of thefront-end detector that detects one gamma may be much better than thetiming or spatial resolution of back-end detector that detects the othergamma of the pair.

The use of front-end and back-end detectors permits flexibility as towhich detector parameters to adapt (temporal, spatial, energyresolution) as well as selected detector material properties (density,Compton scatter versus photoelectric interaction probability, Compton orphotoelectric electron range), for the front-end and back-end detectors.Cost-sensitive decisions can made based on detector characteristics andgeometries in terms of how they influence various PET parametersincluding energy resolution, spatial resolution, temporal resolution,sensitivity, NECR (noise equivalent count rate), true counts,incorrectly classified events, random events, characteristic radiation,Rayleigh scatter, acollinearity, etc.

For example, it may be suitable to employ 0.5 mm thick, high-resistivityor detector grade pixelated silicon or a structured (3D) silicon (orstructured mold) detector arranged edge-on (for adequate energyresolution, improved spatial resolution, faster timing), rather than 1.0mm thick, detector grade Silicon arranged face-on. Or a material with ahigher Z than Silicon could be employed to increase photoelectricinteraction probability (Ge, GaAs CdTe, CZT, structured 3D andstructured mold detectors, etc.). One possibility is that a front-enddetector alone will be adequate. For a dual-layer (or multilayer)detector system all detector interaction combinations (and thus a rangeof PET images with different properties) need to be considered.

Consider a dual-layer detector in which the two layers may have one ormore different properties such as stopping power, spatial resolution andtiming resolution. The first layer could, for example, be comprised ofan array of edge-on, high spatial resolution (typically small pixels),fast or very fast temporal resolution, (low-Z) silicon or structuredsilicon 3D detector planes (or structured mold detectors) providing 3Ddetector capability. The second layer could be an array of edge-on orface-on, moderate or high-Z, semiconductor or scintillator or structured1D or 2D detector planes (providing 2D or 3D detector capability,respectively) of the same or lower spatial and temporal resolution(typically slower, larger pixels).

Photoelectric interactions that occur in the first layer or second layeras well as valid reconstructed events (the result of tracking of singlescatter or multiple scattered photons as well as shared energy within orbetween layers) can be used in coincidence detection with an opposingdual-layer detector. Note that this will result in multiple PET images.Coincidence between opposing (typically faster, smaller pixels) firstlayer detectors (based on photoelectric events or tracked scatter eventsinteracting within the first layer) may be best for spatial and timingresolution information followed by coincidence between a first layerdetector and an opposing second layer detector. The poorest spatial andtiming resolution would be provided by coincidence between (typicallyslower, larger pixel) second layer detectors.

As in the case of a dual-layer or “telescope” Compton camera whichemploys tracking the synergistic interaction of the detector layersenables the recovery of a fraction of the scatter events that interactwithin one or both layers. A low-Z semiconductor material such assilicon, in which Compton scattering dominates the photoelectric effectat 511 keV, can be used in high resolution PET imaging since thephotoelectric and scattering effects can be exploited. If Comptonreconstruction algorithms can be employed the effects of acollinearityand scatter may be reduced in some instances and some non-coincidencedetected events can be used to form a non-coincidence PET image(Compton-PET image). If advantageous, data from one or more types ofacquired PET images can be combined to reconstruct enhanced PET images.Note that low-Z semiconductor detectors such as silicon as well asstructured mold semiconductor detectors and other structuredsemiconductor detectors (for example, structured 3D silicon detectors)are also suitable for use in PET detector systems as well as CT detectorsystems (ring, partial ring, cone beam) in single layer or multilayer oroff-set layer detector formats.

CT-Compton-PET Detector Systems

The flexibility of the Compton camera design allows it to be adapted forPET (and nuclear medicine) imaging. The Compton camera design can alsobe adapted for use in diagnostic x-ray imaging applications such as CTand projection radiography (with the understanding that typical datarate requirements will be much higher, spatial resolution requirementsmay increase, and the operational energy range for diagnostic medical CTis typically lower than for PET and nuclear medicine imaging).

Various coincidence and non-coincidence Compton-PET detector systemimplementations have been described. An extension of this dual-useconcept is to describe a multi-use CT-Compton-PET detector system design(with the understanding that nuclear medicine imaging capability canalso be implemented).

The incorporation of CT features can be explained by examining a specialcase of a Compton-PET detector system design, the CT-Compton-PETdetector system design. This is of interest because CT-PET detectorimaging systems are commercially available. However, the CT and PETdetector imaging sub-systems (which use face-on detectors) arephysically distinct. This commercial configuration involves moving thepatient with respect to the typical partial ring geometry (oralternatively a cone beam geometry) CT scanner into a physicallyseparate PET scanner. These CT and PET detector sub-systems do not sharedetectors or the image acquisition space.

An alternative to the existing commercial CT-PET detector imagingsystems are improved CT-PET detector systems in which the CT scanner orPET scanner (or both) are replaced with novel edge-on CT scanners and/orPET scanners (including Compton-PET detectors) described in thisapplication. For example, the face-on detector CT configuration isreplaced with an edge-on CT detector system capable of performing atleast one of energy integration, PC, and PCE (Nelson, U.S. Pat. No.6,583,420; U.S. Pat. No. 7,291,841; U.S. Pat. No. 7,635,848; U.S. Pat.No. 8,017,906; U.S. Pat. No. 8,115,174; U.S. Pat. No. 8,115,175; andU.S. Pat. No. 8,183,533).

For example, one CT implementation would employ a single layer using anarray of edge-on semiconductor detectors operating in PCE mode (such asan edge-on semiconductor silicon detector or an edge-on structuredsemiconductor detector). Suitable edge-on structured semiconductordetectors include at least one of a 3D semiconductor detector such as 3Dsilicon, a structured mold semiconductor detector incorporating one ormore of semiconductor quantum dots, amorphous semiconductors, andpolycrystalline semiconductors. If additional stopping power is needed asecond layer of moderate-to-high Z semiconductor or scintillatordetectors could be implemented.

One implementation of a second detector layer is to employ face-on oredge-on semiconductor arrays (including structured detectors) whichoperate in PCE mode (based on factors such as cost, detector responselimitations, and/or information content of the radiation field).Less-costly implementations operate in limited PCE mode (two or moreenergy bins) or PE mode. An alternative (less-costly) implementation isfor the second layer to employ a face-on or edge-on scintillator arrayoperating in integration mode or PC mode to primarily detect the moreenergetic x-rays (providing additional information about the radiationfield to compliment the spectral information acquired with the firstdetector layer). If advantageous, detectors can implement mode switchingcircuitry (for example, from PCE mode to integration mode or PC mode;from PE mode to integration mode) as a technique for compensating forexcessive event rates.

A fast, improved CT-PET detector system incorporates multiple x-raytubes (two, three or more) or x-ray sources (such as carbon nanotubes,scanning electron beams, etc.) to reduce image acquisition times. NovelPET detectors include, but are not limited to, 3D crossed rod, crossedfiber-rod and encoded PET detectors. The physically separate PET orCompton-PET scanner preferably provides one or more detector featuressuch as suitable or excellent energy resolution, 3D spatial resolutionand TOF capability. If reduced PET performance is acceptable then one ormore of energy, spatial and temporal resolution can be degraded.

PET designs described in this patent application can be employed withcommercial face-on CT scanners to comprise enhanced CT-PET detectorsystems. Physically separate commercial PET scanners can also be usedwith an edge-on CT detector system in another version of an enhanceddual CT-PET imaging system. Still another version of an enhanced dualCT-PET imaging system employs physically separate edge-on CT and PETdesigns described in this application and prior patents. Yet anotherversion of an enhanced CT-PET imaging system is to employ a face-ondetector or edge-on detector CT scanner with a physically separateCompton-PET detector system.

Alternative to commercial and enhanced dual CT-PET detector designs areCT-Compton-PET systems in which detector components and/or space areshared, representing a cost effective and compact design compared withthe benefit that the patient remains stationary and so registrationbetween CT and PET images is straightforward. Furthermore current CTimaging sub-systems in commercial dual CT-PET systems do not offer PC orPCE capabilities which are available in enhanced dual CT-PET andCT-Compton-PET detector systems.

PC or PCE capabilities can be used for dose reduction and/ormultispectral imaging. Furthermore, multispectral imaging can beimplemented with a PC detector system by implementing x-ray tube voltageswitch (currently employed with dual-energy CT detector systems).

CT-Compton-PET detector systems designs incorporate the capabilitiesdescribed for Compton-PET detector systems. One or more layers ofdetectors can be employed. PET options include non-coincidence(one-sided) and coincidence PET imaging capabilities. The incorporationof x-ray CT capabilities may impose additional requirements on thedesign of the radiation detectors, depending on the energy range for theapplication (small animal, pediatric, adult, therapy, industrial, HLS,synchrotrons) and the event (data) rates (which, for medical CT imaging,are typically much higher than the event rates encountered in nuclearmedicine imaging).

In addition, collimation may be introduced into the CT detector whichwould be of a relatively fine nature. The type and amount of collimationintroduced into the CT detector configuration is preferably sufficientto at least result in a beneficial reduction in radiation cross talkbetween detector elements during CT imaging without substantiallyreducing the efficiency of the PET detector component of the imagingsystem. If external collimation is employed to reduce the intensity ofx-rays scattered by the object from reaching the CT detector, and thisexternal collimation has an undesirable impact on PET imaging efficiencyor image quality, then the external collimation is preferably moveableso that it can rotate or slide out of the detector field of view (FOV)during PET imaging.

X-ray scatter correction algorithms in CT imaging can also be employedwith or without collimation along with corrections for detector effectssuch as induced charges in nearest-neighbor detector elements, chargecloud diffusion and radiation cross talk (energetic electrons,characteristic x-rays, bremsstrahlung) between detector elements(Nelson, U.S. Pat. No. 7,291,841; U.S. Pat. No. 8,017,906). If the PETdetector imaging is not implemented simultaneously with the CT detectorimaging then an optional movable, attenuating shield (such as, but notlimited to Cu, W, Pb, a multilayer material) can be inserted during CTimaging to protect the PET detector from unnecessary radiation damage,and then removed during PET imaging.

The insertion of nuclear medicine collimation hardware such as parallelor pin hole (pinhole) collimators into these Compton camera designs canprovide nuclear medicine imaging capabilities for those cases in whichthe Compton camera does not offer adequate imaging properties. CTdetector modes of operation can include energy integration, PC or PCE.One implementation of a CT-Compton-PET detector system is to simplyoperate the back-end PET detector independently of the front-end CTdetector and accept that the CT detector acts as an attenuator andscatterer of the 511 keV PET gamma rays.

More sophisticated CT-Compton-PET detector systems will be describednext. Implementations of detector geometries include planar (and focusedplanar) configurations and focused structure configurations such asrings and partial rings (as well as focused rings and focused partialrings). Planar, ring, and partial ring detector geometries areencountered in medical diagnostic x-ray CT.

CT-Compton-PET detector system designs described herein are based onimplementations of coincidence and non-coincidence Compton-PET detectorsystems with additional constraints imposed by CT imaging. X-ray fluencerates for diagnostic medical x-ray CT are typically sufficiently highthat features such as PC and PCE are easier to implement if thedistribution of detected events during a time interval is spread outover a number of detector channels. Other constraints on detectorselection are related to problems such as dose-dependent pixelperformance degradation (including polarization issues) and detectoreffects described herein.

This tends to limit the selection of edge-on or face-on detector to oneor more fast-to-very fast, low-to-moderate Z semiconductors (andvariants thereof) with or without gain capability (including, but notlimited to, Si, Ge, GaAs, diamond, Se, Si-APDs, SiPMs, siliconnanowires, iDADs, Se-APDs, GaAsPMs, DiamondPMs), structured 3Dsemiconductor detectors and structured mold (quantum dot, amorphous,polycrystalline) semiconductor detectors coupled to high speed readoutcircuitry (such as a custom readout ASIC or a Medipix chip). Otheroptions include configurations such as gas-based Medipix detectors andfast-to-very fast, bright scintillators coupled to photodetectors.

Other semiconductor material such as CdTe or CZT may be employed if theyare sufficiently thin (typically less than 1 mm) such that issuesrelated to high data rates can be mitigated. Their pixel performancedegradation rates and detector effects should be acceptable (or can, inpart, be compensated by evaluating whether any correlated charge wasdeposited in neighboring pixels (charge sharing) as in the case of theMedipix detector chip). For conventional semiconductor detector designs(including silicon detectors) the importance of implementing correctionsfor charge sharing (as well as x-ray cross talk) typically increases asthe photon energy increases, the pixel size decreases and pixelthickness increases. For edge-on and face-n detector designs chargesharing can occur between neighboring pixels within the same detectorarray and between neighboring detector arrays (for example, adjacentedge-on detector arrays or adjacent face-on detector layers).

For the case of a focused structure detector geometry such as a ring thedetector modules can form partial rings, with detectors in a singlepartial ring that have small gaps or gaps comparable in thickness to 2Dedge-on detector plane modules (with optional collimation between thedetector plane modules). If gaps are of comparable thickness to the 2Dedge-on detector plane modules then the x-ray source is preferablycollimated to match the gaps in the detector plane and the collimatorsand detector need to move along the ring by one pixel width (detectorplane width) to acquire a complete projection for reconstruction. Thiscompensating motion and matching x-ray source collimation is not neededif at least two sets of partially-offset or completely-offset detectorrings (alternate detector modules are located at two different radii)with gaps comparable to the thickness of 2D edge-on detector modules areemployed (Nelson, U.S. Pat. No. 7,291,841).

The CT edge-on detector modules employed in a focused structure ringgeometry can also be employed in a planar CT detector geometry. One ormore layers of edge-on detector modules can be configured to be parallelor tilted with respect to adjacent detector modules in order to achievea focusing effect. As with the ring geometry implementations, layers oftilted edge-on detector modules can also be partially offset orcompletely offset so that tilted edge-on detector modules in a lowerlayer fill gaps between edge-on detector modules in the upper layer(s)so that a reasonably continuous detector is emulated.

As described, a focused pixel structure can be implemented along thelengths of the edge-on tilted (or parallel) detector modules. Variousconfigurations of edge-on or face-on (single or multilayer) detectormodules or combinations of face-on and edge-on detector modules may alsobe employed in planar and ring detector geometries. Optionally, SAR andDOI capabilities can be incorporated into the edge-on and face-ondetector modules, respectively.

If the front-end CT detector and the back-end Compton-PET (or PET)detector operate independently of each other, then the CT-Compton-PETdetector system can be considered a limited CT-Compton-PET detectorsystem (an integrated limited CT-Compton-PET detector system). In thiscase the range of front-end CT detector designs extends from planar tofocused structure (ring and partial ring) geometries and fromtraditional (low-cost) energy integration detectors to PC to PCEdetectors.

The front-end CT detector attenuates a fraction of annihilation gammarays directed toward the back-end Compton-PET (or PET) detector. Theplanar or focused structure back-end Compton-PET (or PET) detector doesnot have to occupy the same FOV as the CT detector; larger or smallerFOVs can be implemented according to hardware constraints, cost anddesired acquisition times. The back-end Compton-PET (or PET) detectorcan be designed to operate with 2D or 3D spatial resolution.

Non-coincidence PET (one-sided PET) imaging can be implemented with alimited CT-Compton-PET system in which the back-end detector is aCompton-PET detector. For coincidence PET imaging, the back-endCompton-PET (or PET) detector can provide either 2D or 3D spatialresolution capability.

Coincidence PET imaging may require the addition of a second PETdetector system and the appropriate coincidence circuitry. If theCompton-PET detector offers 3D resolution and tracking capability thenboth coincidence and non-coincidence PET imaging can be conductedsimultaneously. Another implementation of a limited CT-Compton-PETdetector system is to position the Compton-PET (or PET) detector outsidethe FOV of the CT detector. A radiation shield may be inserted betweenthe CT detector and the Compton-PET (or PET) detector during CToperation to limit unnecessary radiation dose to the Compton-PET (orPET) detector system.

For example, consider the case wherein at least one CT detector array(also referred to as a CT detector) is employed (high speed acquisitionmay require multiple distinct CT detector arrays and x-ray sources oreven a full ring CT detector array). An implementation for a ring-likeacquisition geometry can employ at least one front-end (inner layer) CTdetector as a partial ring detector (or a planar detector) for PET(preferably) aligned with a pair of back-end (outer layer) opposingpartial ring (or planar) PET detectors or a full ring PET detector.

Optionally, a fraction of the CT detector can be implemented with PETdetection features. (For the case of a dedicated PET or Compton-PETimaging system the one or more partial ring (or planar) CT detectorarrays or a full ring CT detector array can be replaced with comparable(or smaller) detector arrays with the advantage that the pixel geometryand readout electronics can be adapted for PET or Compton-PET imagingwithout consideration for CT pixel geometry or CT readout electronics.)Note that in an alternative implementation at least one front-end CTdetector array with a planar format can be aligned with a pair ofback-end opposing planar or partial ring PET detectors or a full ring(or square/rectangular) PET detector. One or more (aligned or unaligned)additional front-end (AFE) detectors (not necessarily used for CT) canbe incorporated into the detector geometry in order to improve PETimaging system capabilities such as detection efficiency and/or timingresolution.

For example, an AFE detector could be paired with the CT detector (or afraction thereof). Extra AFE detectors can be added, either positionedindependently or positioned as opposing pairs. The case of a singlefront-end partial ring CT detector aligned with a pair of back-endopposing partial ring PET detectors permits many interactions to beconsidered (the partial ring CT detector can also interact withpartially-aligned or unaligned PET detectors). A Compton image can beacquired by using the CT detector for scattering gammas, a fraction ofwhich are then preferably detected by either the aligned PET detectorbehind it or by the CT detector alone (e.g., if it has 3D capability).Another option is for the PET detector alone (if it has 3D capability)to scatter and detect incident gammas that did not interact with the CTdetector. The information from Compton scattering a sufficient number oftimes within or between detectors can be used for Compton imagereconstruction based on multiple scattering.

Alternatively, the information from Compton scattering one or more timesterminating with a photoelectric event can be used for Compton imagereconstruction. Compton image reconstruction can also be implemented inthe opposing, aligned PET detector (e.g., if it has 3D capability). If aCompton image event can also be used for PET coincidence imaging thenthe Compton image information may be used to enhance the PET image sinceadditional information concerning the directionality of the detectedgamma is available. PET coincidence images can be acquired based oncoincidence between a PET detector with a PET detector, a CT detector ora CT/PET detector (a CT/PET detector represents, e.g., a CT detectorinteracting with preferably an aligned back-end PET detector althoughthe CT detector can also interact with partially-aligned and unalignedback-end PET detectors).

If at least one AFE detector is present then new coincidence images canbe acquired including AFE detector coincidence with a CT detector (andAFE/PET if the AFE can interact with a PET detector), a PET detector, aCT/PET detector and other AFE (as well as AFE/PET) detectors (ifpresent). Furthermore, AFE (and AFE/PET, if present) Compton images canbe acquired. Image data can be combined when appropriate to synthesizeenhanced diagnostic images. The dual-layer detector formats describedherein can also be implemented as dedicated PET detectors (or nuclearmedicine/PET detectors) including the benefits of generating multipletypes of coincidence images and/or Compton images, using Compton imageinformation to enhance coincidence PET imaging and the synthesis ofenhanced diagnostic images. Note that each layer in the dual-layerdetector formats described can be comprised of sub-layers (and/or have adetector with a structure that is equivalent to having sub-layers).Furthermore, detector properties within a layer or sub-layer can vary (adesign feature also applicable for dedicated multilayer and single layerPET systems).

For example, consider the front-end partial ring CT detector used with aback-end PET ring detector with detectors that offer only moderatetemporal resolution (not necessarily suitable for TOF PET). Thefront-end partial ring CT detector typically offers high spatialresolution but it may or may not offer high temporal (fast or very fast)resolution. If high temporal resolution for the PET image associatedwith the CT detector is desired and the CT detector is not fast then thealigned PET detector segments behind and opposite the CT detector shouldpreferably implement suitably fast detectors (e.g., replacing both ofthe moderate temporal resolution PET detector segments with fast PETdetector segments). Alternatively, if the CT detector is fast then thealigned PET detector segment opposite the CT detector should be suitablyfast whereas the aligned PET detector behind the CT detector can be fastbut it is not required to be fast (if the interaction probability of 511keV gammas with the front-end CT detector is satisfactory for generatingsufficient coincidence events with the opposing PET detector segmentduring image acquisition). Then one or both moderate temporal resolutionPET detector segments may be replaced with suitably fast or very fastdetector segments.

Since the back-end PET detector segments are expected to detect afraction of non-scattered gammas another option is to implement fast PETdetector segments regardless of the temporal response of the front-enddetectors. Yet another option is to introduce at least one AFE detectorthat may or may not offer high temporal resolution to operate with theopposing PET detector segment permitting increased flexibility withrespect to properties implemented for the opposing PET detector segment.The PET detector segments for CT-Compton-PET or dedicated PET systemscan optionally be designed to include positioning capability permittinggreater flexibility for improved image acquisition. Cost can factor intothe decision if the fast PET detector segments cost substantially morethan the PET detector segments they would replace. Other tradeoffs suchas differences in detector stopping power and readout electronicsrequirements need to be considered.

For CT-Compton-PET detector systems the front-end CT detector alsoserves as front-end detector layer for a Compton-PET detector system.The readout electronics should be suitable to handle event data ratesthat are on a comparable scale to the event data rates experienced by CTdetectors, or the CT detector pixel geometry can be modified to reducethe effective data rate per pixel and so reduce the requirements of thereadout electronics. The front-end and back-end detector layerspreferably include appropriate internal and intra-layer event trackingcapabilities (for coincidence and non-coincidence Compton-PET systems)depending on their intended use.

For CT applications which utilize PC or PCE capabilities several edge-onpixel geometries have been described including uniform pixel sizes (1Dor 2D pixel array) and non-uniform pixel sizes (Nelson, U.S. Pat. No.7,635,848; U.S. Pat. No. 8,115,174; U.S. Pat. No. 8,115,175; and U.S.Pat. No. 8,183,533). Issues arise as to x-ray beam hardening with depthof penetration and the benefit of imposing a more-uniform distributionof interaction rates between pixels along the x-ray beam direction(reducing readout errors and readout electronics costs).

If the event rate is sufficiently low a uniform pixel distribution maybe adequate even if beam hardening occurs with penetration depth. If theevent rate is high (as expected in many diagnostic medical x-ray CTapplications) and PC or PCE capability is required, then a static,uniform 2D pixel array may not offer a good balance in detected eventrate per pixel unless the pixel dimensions are relatively small in termsof the stopping power of the detector material. Implementing suchrelatively small pixels allows a degree of flexibility since a variableeffective pixel size versus depth could be synthesized by combining theoutput signals from two or more pixels.

Unfortunately, as pixel size decreases the number of pixels and readoutelectronics increases which raises the cost of the detector modules. Inaddition to detector effects pixel readout noise can increase due toleakage issues associated with some small pixel implementations.

High event rates and x-ray beam hardening with penetration depth mayfavor the use of a non-uniform pixel size with increasing detector depthalong a pixel column. The pixel length within a column can increasedwith increasing depth, resulting in a non-uniform (variable) readoutelement pitch in order to provide a more-balanced count rate per pixelfor the readout electronics.

Detector pixel distributions as well as the use of collimating septaand/or side shielding for detector modules used in CT systems have beendescribed (Nelson, U.S. Pat. No. 6,583,420; U.S. Pat. No. 7,291,841;U.S. Pat. No. 7,635,848; U.S. Pat. No. 8,017,906; U.S. Pat. No.8,115,174; U.S. Pat. No. 8,115,175; and U.S. Pat. No. 8,183,533).Furthermore, the pixel size in the axial direction (the slice thickness)can be non-uniform (benefitting dose reduction). For example, a highresolution pixel size (thin slices) could be implemented near the centerof the detector in the axial direction with a lower resolution pixelsize (thicker slices) implemented on both sides of the center.

Additional non-uniform pixel size distributions can be implemented basedon imaging requirements. Additional flexibility is provided when theoutputs of two or more pixels in the axial direction can be combinedelectronically in order to synthesize the desired distribution of pixelsizes in the axial direction. A non-uniform pixel size in the axialdirection can be implemented with edge-on detectors and/or face-ondetectors. A non-uniform pixel size distribution can be implementedalong an arc segment.

For example, in one implementation of a non-uniform detector pixel sizein the axial direction the high resolution detector pixels could becentered mid-way between the axial limits of the detector arc with lowspatial resolution detectors on either side. In this case the region ofinterest within the object being scanned that benefits from higherspatial resolution is radially-opposite the high resolution detectorpixels whereas the regions of interest on either side (for which lowerspatial resolution is acceptable) are radially-opposite the lowresolution detector pixels. Arrangements of non-uniform detector pixelsdistributions include at least one detector region with high or moderateresolution detector pixels (detector regions with graded spatialresolution can also be employed).

Relative positioning of the high resolution, moderate resolution and lowresolution detector pixels in the axial direction can be adapted for aspecific imaging application. The choice of PCE, PC, and integratingreadout electronics (or combinations thereof) can be adapted for theimaging requirements of individual detector regions. Furthermore,edge-on detectors, face-on detectors or combinations thereof can be usedto implement these arrangements of non-uniform detector pixels.

With edge-on detectors the low spatial resolution detectors can besynthesized by combining the outputs of two or more pixels with the samecoordinates as measured with respect to the edge-on detectorsthemselves. Thus, comparable pixels from adjacent edge-on detectors(even if they are offset with respect to their neighbor) are combined.The same capability to combine the outputs of two or more pixels inorder to reduce spatial resolution can be implemented with face-ondetectors. The ability to synthesize larger pixels dynamically allowsthe edge-on or face-on detector to operate with either uniform or anon-uniform spatial resolution.

An alternative is to implement the edge-on or face-on detector with afixed, non-uniform spatial resolution. The advantage of positioningdetectors with higher spatial resolution closer to the region ofinterest is not limited to CT imaging, this approach to non-uniformdetector spatial resolution can also be implemented for Compton, SPECTand PET imaging (as well as scientific and industrial imaging). Detectornon-uniformity can be implemented along at least one of the axialdirection, the radial direction, an arc. Detector non-uniformity canalso be implemented as an insert to an existing detector system.

Furthermore, this approach to employing detectors with non-uniformproperties such as spatial resolution can be implemented for otherdetector properties including temporal resolution and energy resolution(and even radiation resistance), independent of employing detectors withnon-uniform spatial resolution. One or multiple detector materials canbe employed including scintillators, semiconductors, structureddetectors, etc. and combinations thereof. PCE, PC, and integratingreadout electronics can be employed as well as combinations thereof. Forexample, a CT scanner could implement detector regions in which PCE isemployed whereas other detector regions employ PC or integratingelectronics.

Both PCE and PC readout modes can be deployed as needed according to theimaging requirements along the axial direction and along the arc (suchas the need for energy subtraction in a limited region of image).Alternatively, an integration readout mode can be implemented if a PCEor PC readout mode is or will be saturated (or is not needed).Appropriate beam collimation and filtration can be employed to match thepixel distribution in the axial direction and along the arc.Furthermore, non-uniformity can be extended to include the detectorgeometry type (mixing of edge-on and face-on detectors). For example,high spatial resolution edge-on detectors are (in one implementation)positioned at the middle of the detector arc (that images the region ofinterest within the object being scanned), with low spatial resolutionface-on detectors on either side (potentially reducing the over-all costof the detector system).

The principles of non-uniformity in pixel size (as well as temporalresolution, energy resolution, radiation resistance) and detectorgeometry type can be applied to both ring and planar detector systems.Detector configurations of reduced size can be employed if region ofinterest CT is implemented (retaining the high spatial resolutiondetectors that image the region of interest within the object beingscanned while eliminating the low spatial resolution detectors on eitherside).

A focused structure, ring geometry Compton camera design (Nelson, U.S.Pat. No. 7,291,841), may or may not offer optimal performance as aCT-Compton-PET camera for high event (data) rate, fan beam CT diagnosticimaging. The Compton camera would preferably use edge-on detectormodules with a uniform pixel size along a column (uniform 3D spatialresolution), whereas the PC or PCE CT system would preferably useedge-on detector modules with a variable readout element pitch along acolumn.

The variable readout element pitch for CT allows the readout raterequirements of the readout ASIC-based electronics to be better balancedbetween readout elements (pixels) near the entrance surface and pixelsdistant from the entrance surface, which experience reduced beamintensity. Thus the number of readout elements can be reduced noticeablyand fewer readout ASICs of a given performance level are needed comparedto a uniform pixel array with many small pixels. If the readout ASICselectronics offer high readout data rates sufficient to handle themaximum expected CT data rates for any pixel in a uniform pixel detectorwhich is preferred for use in a Compton camera or Compton-PET detector,then this not an insurmountable constraint.

A potential drawback is a likely increase in cost due to a need for morehigh speed readout ASICs than would be utilized in a dedicate CT scannerwith similar PC or PCE capabilities, but a non-uniform pixeldistribution with depth. Other issues that may arise due to thisCT-COMPTON-PET detector system design and the increased use of highspeed readout ASICs are related to an increase in heat generation andtherefore new cooling requirements to avoid increased detector noise andthermal expansion issues. There is also a possibility that some readoutASICs may be moved closer to the pixels (which may result in certainreadout ASICs positioned within the x-ray beam path and thereforealtering shielding requirements).

Note that this issue of CT detectors with uniform and non-uniform pixelarrays in CT-Compton-PET detector systems affects both the focusedstructure ring (or partial ring) detector format used in fan beam CT andthe planar detector format used in cone beam CT. One alternative is touse readout ASICs of varying performance with respect to depth. Thehighest speed readout ASICs would read out the pixels close to theentrance surface, whereas readout ASICs of progressively slower speeds(but still sufficient for both CT and Compton camera applications) couldbe used to read out pixels at greater depths.

Another alternative is to enable the edge-on detector module electronicsto redefine the readout element pitch according to whether theCT-Compton-PET detector system is functioning as a PET detector systemor a CT detector system. Thus, a detector module can have a selectable(fixed or variable) effective pixel width along a detector row and/or aneffective pixel length along a detector column in which the effectivepixel width or length is synthesized from the outputs of one or more(typically) uniformly spaced pixels.

For example, a variable, effective pixel length can be selected for CTimaging based on the beam spectrum and the beam intensity. A softerx-ray beam would preferentially be attenuated closer to the detectorentrance surface than a harder x-ray beam, for a given detector material(for energies away from a detector material k-edge). For the case of asofter x-ray beam of a given intensity the balancing of event ratesbetween successive effective pixels in a column would benefit fromelectronically synthesizing relatively smaller effective pixel lengthsnear the entrance surface. Relatively larger effective pixel lengthswould create a better balance of event rates between effective pixels inthe case of a harder x-ray beam of a given intensity.

The advantage of a synthesized readout is that it can be adaptedaccording to the energy spectrum and the desired readout rates, thusexpanding the use of a PC or PCE CT system to a broad range of beamspectrums (applications) while retaining the uniform detector pixelgeometry useful for PET (and Compton camera) imaging. Since a SPECTcamera employs collimation to define directionality of the incidentphotons, either uniform or non-uniform detector pixel geometry can beemployed (making a CT-SPECT detector system relatively straightforwardto implement with appropriate collimation in place).

If tracking of Compton-scattered photons within the SPECT camera isimplemented, then a uniform detector pixel geometry may be beneficial.Features such as redefining the readout element pitch (synthesizing aneffective pixel length or width) or employing readout ASICs of varyingperformance with detector depth can be implemented in dedicated CTdetector systems, as well as CT-Compton-PET detector systems andCT-SPECT detector systems. Furthermore, CT-SPECT detector systems canemploy a single detector layer or multiple detector layers.

CT-Compton-PET detector system geometries include planar and focusedplanar detector systems and focused structure detector systems such asring and partial ring (as well as focused ring and focused partial ring)detector systems. Non-coincidence and coincidence CT-Compton-PETconfigurations are described herein based on non-coincidence andcoincidence Compton-PET configurations. The CT x-ray detectors offersuitable 3D spatial resolution, energy resolution (PCE capability) andtemporal resolution to be useful for the high x-ray fluence ratesencountered in medical and non-medical CT scanning as well as for use asthe front-end detector in non-coincidence and coincidence Compton-PETdetector systems. Event tracking capability may be required forCT-Compton-PET systems.

Non-coincidence CT-Compton-PET detector systems combine CT imagingcapability with one-sided PET imaging capability by employing the CTx-ray detector as the front-end detector layer that would be used in anon-coincidence Compton-PET detector system in conjunction with ahigh-stopping power back-end detector. A flexible design employsfront-end and back-end detectors that offer suitable 3D spatialresolution, energy resolution and temporal resolution.

Both the front-end and back-end detector layers provide adequatetemporal resolution for an expected event rate, such that accurate eventtracking can be enabled both within the front-end and back-end detectorsand between the front-end and back-end detectors, since Compton scatterand photoelectric interactions can be recorded in both front-end andback-end detectors. All implementations described for non-coincidenceCompton-PET (three two-layer Compton cameras and three Compton telescopecameras) are applicable, possibly with the added constraint that thefront-end detector should offer suitable detection efficiency for thex-ray energy spectrums that would be used in CT imaging, should becompatible with the event rates for CT imaging, and should offer aspatial resolution with depth that is reasonably uniform when Comptonand/or PET imaging modalities are employed.

FIG. 5 illustrates a perspective of a CT-Compton-PET detector system1000 in a focused structure (partial ring) geometry which includes apoint-like x-ray 109 radiation source 125 and a gamma ray 107 radiationsource 111. The front-end detector layer 510, comprised of detectormodules 102 which use 2D pixelated array radiation detectors 115 in anedge-on geometry with base 106 and communication links 103, performs thedual role as an x-ray CT detector and a front-end detector layer(detector layer 1) for a Compton-PET detector system.

The detector modules 102 are mounted in a rigid structure 110. Theback-end detector layer 520 (detector layer 2) could be of a planar orfocused structure geometry. For comparison, FIG. 1 can be understood toshow the front-end and back-end detector layers 510 and 520 (detectorlayers 1 and 2) for a planar CT-Compton-Pet detector geometry if thefront-end detector layer 510 is suitable for CT imaging.

A reduction in cost can be realized if the Compton-PET capability isimplemented only within a sub-region of the CT detector (for example, asegment of a partial ring detector geometry or a region of a planardetector geometry). In these instances segments of CT detector modulesor regions of CT detector modules that are not involved in PET imagingdo not need to implement features such as synthesizing variableeffective pixel lengths or employing readout ASICs of varyingperformance with detector depth. Multiple Compton-PET views can still beacquired as a result of detector rotation (in some applications theobject can rotate and the detector is stationary).

By reducing the active detector area the detection efficiency will bereduced and acquisition times will, in general, increase. Alternatively,acquisition times can be typically be reduced by increasing the PETdetector FOV beyond the CT detector FOV. As described, if the Comptoncamera image quality isn't suitable for the nuclear medicine imagingapplications of interest then a collimator can be inserted in front ofthe detector so that the system of collimator and detector can functionas a SPECT/gamma camera.

Coincidence CT-Compton-PET detector systems extend the implementationsof non-coincidence CT-Compton-PET detector systems with the addition ofcoincidence detection capability by introducing a second Compton-PETdetector system along with appropriate coincidence circuitry. If theCompton scatter capability of a front-end detector is not needed thenonly a PET-compatible detector may be needed for the second detectorsystem.

Implementations described for coincidence Compton-PET detector systemsare applicable. Thus, the detector geometries shown in FIG. 1 and FIG. 5are applicable when employed in a coincidence detection configurationsuch as FIG. 4. Again, a reduction in cost can be realized if thecoincidence Compton-PET or coincidence PET capability is implementedonly within a sub-region of the CT detector (for example, a segment of apartial ring or complete ring detector geometry or a region of a planardetector geometry) and a matching Compton-PET or a PET-compatibleback-end detector of comparable dimensions is positioned opposite thatsegment or region of the CT detector.

Additional cost savings may be realized if the second coincidenceCompton-PET detector system employs a front-end detector that offerscomparable performance to the CT detector when used as part of aCompton-PET detector system, but lacks the extreme performancecapability of a CT detector. Acquisition times can be typically bereduced by increasing the PET detector FOV beyond the CT detector FOV.

Multiple Compton-PET or PET views of a limited volume of the subject canbe acquired as a result of detector rotation about the subject. In someapplications the subject can rotate and the detector is stationary. Byreducing the active detector area detection efficiency may be reducedand acquisition times may increase. If the Compton camera image qualityisn't suitable for the nuclear medicine imaging applications ofinterest, then a collimator can be inserted so that the detector canfunction as a SPECT/gamma camera.

The CT-COMPTON-PET scanner assigns the CT detector to the role of afront-end detector in a Compton-PET detector system when Compton cameraor PET (or nuclear medicine) imaging is implemented. In animplementation of a coincidence Compton-PET detector, the front-enddetector primarily acted as a Compton scatterer (with photoelectricdetection capability) and the back-end detector provided stopping power,energy resolution and temporal resolution sufficient for event trackingwith respect to the front-end detector. Options described for thefront-end detector include, but are not limited to, sufficiently thinplanar semiconductor detectors, structured 3D semiconductor detectors,structured mold quantum dot semiconductor detectors, detectors with SARor DOI capability, low/moderate-Z scintillator detectors and structuredlow/moderate-Z straw detectors (which typically require lower data ratesthan the semiconductor-based detectors).

Furthermore, the front-end CT detector may be a multilayer detector, asdescribed herein. For example, one implementation employs a front-enddetector comprised of a first layer of an edge-on semiconductor followedby an edge-on or face-on scintillator second layer followed by aback-end PET detector which now functions as a third layer (a variationof this detector design employs a suitably-designed back-end PETdetector to function in the role of the second layer of the front-end CTdetector). (Similar detector implementations may be used for theback-end detector although the detector properties may differ betweenfront-end and back-end detectors.)

If the front-end detector offers an acceptable Compton interactionprobability with annihilation gammas and it is fast enough to providethe required coincidence timing resolution (or very fast coincidencetiming if TOF PET imaging is desired), then the back-end PET detectorrequirements can be simplified since its role is primarily to detect(typically through the photoelectric effect) Compton-scattered gammasfrom the front-end detector. If the back-end detector is required toprovide coincidence resolution (including TOF resolution if desired),then the selection of suitable detector materials and detector designsmay be reduced. In one implementation the back-end PET detector is alsoused to detect annihilation gammas that don't interact with thefront-end detector (requiring coincidence resolution or TOF resolution).(Front-end and back-end detectors can function independently as PETcoincidence detectors, front-end and back-end detectors can functioncooperatively as a PET coincidence detector, and front-end and back-enddetectors can function cooperatively as a Compton gamma camera.)

Reduced spatial resolution could be acceptable for a back-end detector(although Compton camera reconstruction accuracy will be reduced orlost) used in coincidence PET imaging, if the front-end detectorprovides adequate 3D Compton-scatter information includingmoderate-to-high spatial resolution. For example, a single detectorblock, a 1D or a 2D detector array could be implemented based on factorssuch as expected count rate, required energy resolution, cost anddesired flexibility.

In general, for both coincidence and non-coincidence Compton-PETdetector systems, a combination of a 3D back-end detector with a 3Dfront-end detector could improve overall detection efficiency. In thisimplementation the back-end detector could detect Compton scatteredphotons from the front-end detector, non-scattered (primary) photonsusing the photoelectric effect, and photons scattered within theback-end detector itself. Face-on and edge-on (or angled) detectordesigns can be employed for the back-end detector as well as thefront-end detector.

PET scan times can be improved by employing additional partial-ring orplanar PET or Compton-PET detector systems that operate with or areindependent of the coincidence or non-coincidence CT-Compton-PETdetector system. These systems are referred to as enhanced coincidenceor non-coincidence CT-Compton-PET detector systems. The amount ofrotation about the object to acquire a more-complete PET image can bereduced.

Another option is to implement a coincidence CT-Compton-PET detectorsystem based on a multiple (two or more) x-ray tube or x-ray sourcesystem. For example, the angular arc of a commercial, dual x-ray tube CTpartial ring detector is approximately twice that of a single x-ray tubesystem. Multiple cone beam imaging can be implemented if there are twoor more x-ray tubes or x-ray sources and corresponding single layer ormultiple layer (multiple energy resolution) planar detectors. (Anexample of a multi-planar detector/x-ray tube CT system developed forhigh speed cardiac and lung CT was the Mayo Dynamic SpatialReconstructor or DSR first implemented in the late 1970s.) Note that ifinterior tomography techniques can be implemented, then x-rayintensities and/or areas of planar detectors (depending on theapplication) may be reduced (Yu, H. and Wang, G., Phys. Med. Biol., Vol.54(9): pp. 2791-2805 (2009)).

For the case of the focused structure partial ring geometry the CTpartial ring detector (the front-end detector) used in a dual x-ray tubeconfiguration can be split into two equal CT partial ring detectorsections so that at least one CT partial ring detector section (and itsback-end detector) can be rotated through 180 degrees when coincidencePET scanning is initiated. This could be particularly beneficial forapplications such as fast scan Cardiac CT in conjunction with CardiacPET CT. Other applications that could benefit from high resolution CTand PET or SPECT (nuclear medicine) imaging capabilities of this systeminclude head imaging and small animal imaging. Note that the back-enddetector might cover only a segment of a CT partial ring or completering detector (or a region for a planar detector). If coincidenceCT-Compton-PET system is implemented the second planar or partial ringPET detector may only need to be comparable in size to the actual PETdetector implemented with the first CT planar or partial ring detector.

The efficiency of a PET detector system can be improved by addingadditional front-end detectors (and corresponding back-end detectors)opposite to, adjacent to or separate from the CT partial ring detectoror the CT planar detector. These front-end detectors could utilize lessdemanding readout electronics and may not require features such as pixelsynthesis since they would only be used for PET imaging and not CTimaging. Note that for the various PET implementations in which anopposing PET detector would block the x-ray beam path the opposing PETdetector is either rotated out of the beam path (the x-ray tube or x-raysource may be physically retracted when not in use) or a small openingis made in the opposing PET detector to pass the collimated x-ray beam(the PET detector rotates with the x-ray tube or x-ray source). Ifphysical gaps are present within the PET detector due to the presence ofone-or-more x-ray sources then the gaps can be filled by removable PETdetector modules or the PET detector can be rotated to sample themissing PET detector volumes.

Multiple x-ray tubes or x-ray sources (as described for fast, improvedCT-PET detector imaging systems) can be employed with enhancedintegrated non-coincidence or enhanced coincidence CT-Compton-PETdetector imaging systems and enhanced limited integrated CT-Compton-PETdetector imaging systems. Both stationary and rotating x-raytube-detector systems can be implemented (both designs have been usedwith dedicated CT imaging systems).

Dedicated (stand-alone) CT detector imaging system in a ring or planardetector geometry can be implemented by reducing the functionality ofthe CT-Compton-PET detector imaging systems described herein. Asdetailed, detectors with fixed (or variable) uniform or non-uniformpixels can be implemented with the requirement that the detectors canperform efficiently at the event count rates per pixel encountered inmedical CT imaging.

CT detectors include single layer and multilayer detectors comprised offace-on detectors and/or edge-on detectors including gas, scintillator,semiconductor, low temperature (such as Ge and superconductor) andstructured detectors (such as structured 3D semiconductor, structuredmold quantum dot and scintillator-photodetector structured detectors).Single layer and multilayer detector designs of Compton camerasdescribed herein can be implemented in a dedicated CT detector imagingsystem with PCE capability (a simplification would be a design thatprovides PC capability). Multilayer designs typically maintain orincrease the atomic number of the detector material for progressivelydeeper detector layers with respect to the radiation entrance surface.

Consider a single layer, edge-on detector implementation for a medicalCT imaging system in which detector planes are aligned with the Z-axisin a ring geometry. 2D Si edge-on detectors with a wafer thickness of(for example) approximately 500 microns (μm) as currently implementedmay be preferred over relatively thick, expensive, face-on CdTe or CZTdetectors in terms of operational lifetime and temporal response.Alternative edge-on detectors of comparable thickness (approximately 500microns) which can offer improvements with respect to the stopping powerand/or temporal response performance of 2D Si at reduced cost comparedto the relatively thick, face-on CdTe or CZT detectors include, but arenot limited to, 2D ZnO (which is also a fast, relatively low-Zscintillator making ZnO attractive as a semiconductor or scintillatordetector material for TOF PET), 2D GaAs, 2D Ge, 2D CdTe and 2D CZTdetectors (including low noise implementations, implementations withgain) and structured detectors (structured 3D semiconductor detectorssuch as 3D Si, 3D GaAs, 3D CdTe, 3D CZT, 3D Ge, etc., as well asstructured mold (amorphous, polycrystalline, quantum dot/nanoparticle)semiconductor detectors.

If cost is an issue and reduced capabilities (such as reduced energyresolution) are acceptable then a structured mold scintillator quantumdot detector could be employed (for example, functioning as an energyintegrator detector for single or multiple energy CT). Since quantum dotdensity can be varied with position and a single quantum dot material ormultiple quantum dot materials can be employed as a function of positionit is readily apparent that the equivalent of a multilayer detector canbe synthesized within a single structured mold quantum dot detector byvarying quantum dot density and/or material as a function of position inregular or irregular patterns, as described below.

This concept of varying material (and/or material density) is readilyextended to structured mold amorphous and polycrystalline semiconductordetectors, structured mold scintillator detectors, etc. This design canbe further generalized to structured mold semiconductor detectors thatinclude two or more of quantum dots, amorphous semiconductors, andpolycrystalline semiconductors, as well as structured mold detectorsthat incorporate (for example) semiconductors and scintillators,semiconductors with gases, scintillators with gases, etc.

In addition, detectors with thickness greater than or less than 500microns can be implemented depending on the image resolutionrequirements for the CT detector imaging system (medical diagnostic,dental panoramic imaging, radiation therapy, industrial, Homelandsecurity, etc.). This single layer, edge-on detector CT imaging systemcan be employed as a single layer PET imaging system and/or a Comptoncamera/nuclear medicine imaging system.

As described, multiple Compton-PET implementations are possible.Furthermore, PET and Compton camera/nuclear medicine imaging can beconducted simultaneously. Depending on the fraction of the ringcircumference covered by edge-on detectors, additional detectors (of thesame or different design) may need to be added to increase coincidencedetection efficiency.

For the relatively small (hardware) pixel sizes employed in currentmedical CT imaging systems, Si is a reasonably efficient detector forthe lower x-ray energies encountered in mammography CT and pediatric CT.For adult CT the efficiency of Si may suffer, particularly for x-rayenergies above (approximately) 40 keV. A compromise, multilayer detectorconfiguration (for example) could employ edge-on, 2D semiconductor orstructured semiconductor detectors (such as 3D Si or GaAs detectors or astructured mold semiconductor detectors with materials such assemiconductor quantum dots, amorphous semiconductors, polycrystallinesemiconductors) as the low-Z or moderate-Z front-end detector, withmoderate-Z or high-Z, edge-on or face-on, back-end detector. (Note thatif temperature requirements can be met then Ge is a candidate as amoderate-Z, face-on or edge-on detector.) The back-end detector, withappropriate capabilities, may not only improve the overall CT detectorperformance but also may be suitable for a different imaging modalitysuch as PET (as described herein). A simplification is to implement asingle layer or offset detector layer format for CT or PET using atleast one of edge-on, 2D semiconductor detectors, structuredsemiconductor detectors (such as 3D semiconductor detectors orstructured mold semiconductor detectors), structured scintillatordetectors.

Consider the case of a multilayer (in this case, a dual-layer) detectorconfiguration in which an edge-on, 2D Si front-end detector(alternatives such as 3D silicon, etc. may also be employed) is employedas the first detector layer. It would be of reduced height compared to asingle layer, edge-on, 2D Si detector implementation and thus lessexpensive as well as reducing the pixel count and limiting the maximumpixel size. The back-end, second detector layer (edge-on or face-on) istypically comprised of a moderate-Z material (including semiconductorssuch as GaAs or CdTe or CZT, scintillators and/or structured detectors),or a high-Z material (including semiconductors, scintillators, and/orstructured detectors) which would emphasize photoelectric interactionswith the high energy photons that penetrate the front-end detector.

One or more types of back-end, face-on detectors can be configured as 1Ddetectors that are positioned beneath each of the 2D Si edge-ondetectors. The thicknesses of appropriate face-on detectors should notbe so great that detrimental effects such as polarization or lightlosses (for scintillators) cannot be mitigated. The cost ofmanufacturing such 1D detectors (material yields, butting pixels,bonding to readout electronics) should be reduced relative to 2Ddetectors. More than one layer of 1D, face-on detectors can be employedand layers can consist of the same or different materials. Furthermore,if enhanced detector performance is desired (one or more of: higherspatial resolution, higher temporal resolution, higher energyresolution) in the back-end detector, a 2D face-on detector can beimplemented even if SAR is not implemented in the front-end edge-ondetector (as noted previously herein, the detector layers need not havethe same spatial, temporal or energy resolution). If 3D information isdesired then DOI capability can be introduced or additional layers canbe added. Note that in some applications (e.g., due to issues that maybe related to cost, temporal resolution, spatial resolution, energyresolution) it may be desirable to employ face-on 1D or 2D (or 3D)detectors in the front-end and edge-on detectors (1D, 2D or 3D) in theback-end. The useful information that can be extracted from theradiation detected within each layer (as well as cost) will determinewhether individual detectors operate as energy integrators, PCs or PCEs.

An alternative is to position a back-end, edge-on 1D or 2D detector(including structured 3D semiconductor and structured mold semiconductorquantum dot, amorphous semiconductor and polycrystalline semiconductordetector implementations) below each front-end, 2D Si edge-on detector.The edge-on, 1D detector is less-costly to manufacture whereas the 2Darray will typically handle higher data rates and offer better energyresolution. This dual-layer CT design could be used for both low energyand high energy imaging applications. Any combination of suitableedge-on detectors including 2D detectors, structured 3D semiconductordetectors and structured mold semiconductor detectors such assemiconductor quantum dot detectors can be employed for the front-endand back-end detectors. In this case the semiconductor quantum dotsfunction as semiconductor detector materials and therefore a structuredmold semiconductor quantum dot detector can also be described as aspecific implementation of a structured mold semiconductor detector.

It should be noted (as described previously herein) that a single layerimplementation based only on an edge-on structured 3D semiconductordetector or a structured mold semiconductor detector of a single low,moderate or high-Z material may be implemented in place of a dual-layerCT design. Structured mold semiconductor quantum dot detectors offeradditional flexibility, beyond simply varying the dimension of the pixelversus depth to control count rates, in that the density of quantum dotscan be varied from low-to-high for individual pixels that constitute thestructured quantum dot detector. Another technique to vary quantum dotdensity is to vary the number of holes within a pixel that are filledwith a quantum dot material. For example, pixels could be of a uniformdimensions while energy-dependent attenuation could increase with depthby increasing the density of quantum dots with depth. Thus, the quantumdot (detector) material can be varied as a function of depth in anedge-on orientation. Similar functionality can be implemented withstructured mold amorphous and polycrystalline semiconductor detectors.As described herein, structured mold detectors can incorporate more thanone type of semiconductor material as well as mixtures of detectormaterial types (semiconductor, scintillator, gas, superconductor).

Furthermore, the selection of back-end detector materials is not limitedto semiconductors or structured detectors. The back-end detectors can beface-on or (1D or 2D) edge-on scintillator detectors (Nelson, U.S. Pat.No. 7,291,841). (Additional face-on detector layers can be implementedas needed and different scintillator materials can be employed indifferent layers.) In general, the back-end detectors can operate as PC,PCE or integrating detectors depending on the application. For example,a PC or integrating back-end scintillator detector can be paired with aPCE front-end detector simplifying detector design. The filtered beamspectrum reaching the back-end detector can be estimated.

In addition, the flexible design permits either or both front-end andback-end detectors to be scintillator-photodetector detectors. Forexample, the (first layer) front-end detector could be a low-to-moderateZ scintillator-photodetector detector with a (second layer) back-endmoderate-to-high Z structured mold quantum dot detector. As noted, thedetectors in each layer can operate as either PC or PCE or energyintegrating detectors. Thus, one implementation would use a low-Z Sidetector with PCE capability for the first layer with a high-Zscintillator-photodetector detector with energy integration capabilityfor the second layer. Furthermore, the flexible design described fordual-layer detector systems can be implemented for multilayer detectorsystem with three or more layers. The single layer and multilayerdetector systems described herein can incorporate one or morenon-detector materials including attenuating materials, scatteringmaterials and conversion materials depending on the interactingradiation field (e.g., particle types, energies).

It should be noted that a single layer implementation based only on anedge-on structured 3D semiconductor detector or a structured moldsemiconductor detector (implementing at least one of semiconductorquantum dots, amorphous semiconductors, and polycrystallinesemiconductors) of a single low, moderate or high-Z material may beimplemented in place of a dual-layer CT design. Structured moldsemiconductor detectors offer additional flexibility, beyond simplyvarying the dimension of the pixel versus depth to control count rates,in that the density of semiconductor materials such as quantum dots(nanoparticles) can be varied from low-to-high for individual pixelsthat constitute the structured mold semiconductor detector. For example,pixels could be of a uniform dimensions while energy-dependentattenuation could increase with depth by increasing the density ofsemiconductor quantum dots with depth.

Structured Mold Detectors

Structured mold semiconductor quantum dot (nanoparticle) detectors maydeploy a single semiconductor quantum dot (nanoparticle) material. Theuse of edge-on, structured mold semiconductor quantum dot (nanoparticle)detectors creates an opportunity to implement a more flexible detectordesign. For example, multiple semiconductor quantum dot materials canalso be deployed such that low-Z/moderate-Z semiconductor quantum dotmaterials are positioned near the radiation entrance surface andmoderate-Z/high-Z semiconductor quantum dot materials are positionedfurther from the radiation entrance surface (a multilayer structuredmold quantum dot detector). Thus, the selection of semiconductor quantumdot (nanoparticle) materials can be selected for different energy rangesand the count rate per pixel as a function of distance from theradiation entrance surface can be more-balanced. Furthermore, as noted,the densities of each of the multiple semiconductor quantum dot(nanoparticle) materials can be varied for individual pixels fromlow-to-high for purposes of improved operation and imaging. Structuredmold semiconductor detectors may also be referred to as structuredsemiconductor conductive mold detectors or semiconductor conductive molddetectors.

The semiconductor quantum dot (nanoparticle) materials can bedistributed in appropriate patterns for the incident radiation fieldutilized for imaging and the modified radiation field within thedetector (examples include Compton cameras, spectral CT, etc.). Forexample, a geometric pattern such as a series of partial concentricrings can be employed to create a focused edge-on detector (with theability to vary semiconductor quantum dot material and density within aring and between rings). Furthermore, the partial concentric rings canbe comprised of offset pixels (gaps between neighboring pixels in apartial concentric ring that are covered by offset pixels in aneighboring partial concentric ring) rather than a continuum of pixels(FIG. 3 demonstrates a similar design wherein the gaps between offsetpixels in the upper edge-on semiconductor detector layer are covered bythe offset pixels in the lower edge-on semiconductor detector layer).

Other semiconductor detector materials than quantum dots (nanoparticles)can also be employed to fill a 1D or 2D array of channels or a 2D arrayof holes of a conductive mold material, including, but not limited to,polycrystalline and amorphous semiconductor detector materials (e.g.,silicon is frequently used as a conductive mold material, in the form ofporous silicon or micromachined silicon). When appropriate, bothchannels and holes can be present in the conductive mold material. In anedge-on orientation the detector material density distribution can beadjusted as a function of depth if desired. For example, a non-uniformdetector material density distribution can be implemented to compensatefor beam hardening with depth for an edge-on CT detector. In order toattain reasonable signal collection efficiency the practical crosssection dimensions for holes (diameters for the case of circular holes)or practical channel widths are limited in part by the travel ranges ofthe information carriers. Typically, as a cross section dimension (ordimensions) increase the difficulty of chemical etching ormicromachining of holes and channels of a desired depth is reduced aswell as the difficulty of attaining an acceptable degree of uniformfilling of holes or channels with semiconductor detector materials.

Adjusting the density distribution of a semiconductor detector materialwithin the detector can be implemented in conjunction with pixel sizeadjustments to provide even greater flexibility in detector adaptationfor an incident radiation field. Thus, semiconductor detector materials(including semiconductor quantum dot (nanoparticle) materials) can bedistributed according to appropriate patterns for the properties of theincident radiation field (types of particles, energies, angulardistribution, intensity distribution, etc.) utilized for imaging. Aspecific implementation (but not the only possible implementation) ofSAR involves segmenting the holes or channels into at least two partssuch that separate signals can be read out from the segments.

Structured mold detectors that employ a single detector material (suchas a single quantum dot (nanoparticle) semiconductor material, a singlequantum dot (nanoparticle) scintillator material, etc.) are basicstructured mold detectors. Hybrid structured mold detectors can becomprised of multiple detector materials of a single type (for example,multiple semiconductor materials or multiple scintillator materials) ormultiple types of detector materials (for example, one or moresemiconductor materials combined with one or more scintillator materialsand/or gas materials, etc.). An example is a hybrid structured moldsemiconductor detector that is comprised of a mixture of semiconductorquantum dot materials and/or other semiconductor materials such aspolycrystalline and/or amorphous semiconductor detector materialswherein low-to-moderate Z materials are positioned to intercept thex-ray beam in the front-end while moderate-to-high Z materials arepositioned in the back-end of the edge-on detector. Furthermore, hybridstructured mold detectors can also incorporate structured 3D detectorssuch that a region of the structured mold detector is nonporous (thestructured 3D detector region) and another region is porous (with poresincorporating detector materials). For example, low-Z 3D silicon ispositioned in the nonporous front-end of the edge-on detector andmoderate-to-high Z semiconductor materials are positioned in the porousback-end of the edge-on detector.

Furthermore, basic structured mold detectors and hybrid structured molddetectors can incorporate one or more non-detector materials includingattenuating materials, scattering materials and conversion materials toprovide transverse filtering and/or lateral shielding. Thus, basicstructured mold semiconductor detectors and hybrid structured moldsemiconductor detectors can incorporate attenuating materials,scattering materials and conversion materials. For example, patterns ofholes or channels can be filled with materials that contain (e.g., forthe case of x-ray and gamma ray radiation) moderate-to-high Zattenuating elements (iron, copper, lead, tungsten, uranium, gold, etc.)or alloys in order to create collimation for the radiation fieldincident on the detector, to reduce inter-detector cross talk elements(lateral shielding), to incorporate spectral (and/or particle type)filters such as gratings (diffraction, absorption, phase). Scatteringand conversion materials can be employed in order to alter theproperties of the incident radiation field prior to being detected(transverse filtering). For example, conversion materials can be used totransform x-rays to electrons, neutrons to photons, fast neutrons tothermal neutrons, etc. for the benefit of the detector. Scatteringmaterials could be used, for example, to degrade photons or neutronswith undesirable energies. The densities of selected attenuating,scattering and converting materials can also be varied with position inorder to improve detector performance.

Although the advantages of basic structured mold (and hybrid structuredmold) semiconductor detectors have been described for an edge-ondetector orientation it should be apparent that similar advantages canbe realized for basic structured mold detectors and hybrid structuredmold detectors employed in a face-on orientation (or tiltedorientation). For example, semiconductor detector materials can bedistributed in appropriate patterns for the incident radiation fieldutilized for imaging, and semiconductor detector material densities canbe varied as needed within a single face-on detector layer or withmultiple face-on detector layers. Furthermore, patterns of holes orchannels can be filled with attenuating materials, conversionsmaterials, spectral (and/or particle type) filters such as gratings(diffraction, absorption, phase), etc. in a face-on basic structureddetector or a face-on hybrid structured detector.

For example, one implementation would create an attenuating grid patternproviding lateral (side) shielding between face-on detector elements.Non-detector materials such as attenuators, scatterers and converterscan also be applied in appropriate patterns on the surface of theface-on basic structured detector or hybrid structured detector (thisrepresents an alternative to introducing a structured layer that onlyimplements appropriate patterns of non-detector materials prior to astructured detector layer). Note that an edge-on SAR capability becomesa face-on DOI capability.

The benefits of basic structured detectors and hybrid structureddetectors are not limited to the implementation of semiconductordetector materials, these benefits are also available forimplementations wherein scintillator detector materials (or otherdetector materials such as gas and low temperature detectors) areemployed. Thus, scintillator materials can be arranged in patterns,scintillator density can be varied, multiple scintillator materials canbe employed, scintillator and non-scintillator detectors can be employedin the same edge-on or face-on basic or hybrid structured detector andnon-detector materials can be employed.

The positioning of detector materials and density distribution ofdetector materials (as well as non-detector materials) within edge-onand face-on basic structured detectors and hybrid structured detectorscan be adapted according to the properties of the radiation field beingdetected (including mixed radiation fields). In a layered detectorsystem either edge-on detectors or face-on detectors or both edge-on andface-on detectors (as well as tilted detectors) can be employed.Furthermore, both structured detectors and more conventional detectorscan be employed within a layered detector system. For example, in a twolayer system an edge-on, basic structured semiconductor quantum dotdetector could be aligned with and followed by a 1D scintillator arraycoupled to a photodetector (or amplified photodetector) array.

Structured mold semiconductor and scintillator quantum dot(nanoparticle) detectors (as well as structured 3D detectors and 2Dsemiconductor detectors) can be implemented with fixed or adjustablepixel sizes which can be uniform or non-uniform. Furthermore, thedensity of quantum dot material can be varied with position. Typicallythe lowest density of quantum dot material could be positioned near theradiation entrance surface.

A moderate-Z or high-Z structured mold semiconductor quantum dot(nanoparticle) detector can also be employed in a face-on orientation asa 1D (or 2D) detector positioned after a (for example) low-Z, 2D Siedge-on detector. Furthermore, moderate-Z or high-Z, fast, brightscintillator-photodetector 1D (or 2D) array detectors (includingstructured scintillator and nanophosphor detectors), face-on or edge-on,can be employed after a (for example) low-Z, 2D Si edge-on detector(providing limited energy resolution or simply providing photon countingor integration capability or acting as an energy integrators).

The photodetector is a fast, sensitive 1D photodetector that can bechosen from (but is not limited to) photodiodes and amplifiedphotodetectors including, but not limited to, APDs, SiPMs, siliconnanowires, GaAsPMs, DiamondPMs, electron multiplier CCDs andmicrochannel plates with a pixel structure or a dual-readout structure.Scintillator-photodetector detectors can employ scintillator screens,deposited scintillator films, ceramic scintillators and cut scintillatorsheets (including thick sheets which can also referred to as blocks orslabs). Scintillator-photodetector structured detectors can employstructured scintillators (such as manufactured scintillator arrays,scintillators that demonstrate columnar growth and scintillators coupledto fiber arrays) as well as scintillating or minifying scintillating,focused or unfocused, fiber arrays.

Scintillating fiber materials include, but are not limited to,phosphors, granular phosphors, nanophosphors and scintillator quantumdots. If limited energy resolution is acceptable or only photon countingis needed for CT then a moderate-Z or high-Z, fast, bright,scintillator-photodetector or scintillator-photodetector structureddetector (wherein the photodetector is a fast, sensitive photodetector)can be used in place of the single layer or dual-layer detectorimplementations as described herein (see Nelson et al., U.S. Pat. No.4,560,882; U.S. Pat. No. 5,258,145; U.S. Pat. No. 8,017,906; and U.S.Pat. No. 9,347,893).

FIG. 6 illustrates a minifying scintillating fiber array 140 coupled toa 1D photodetector 141 which is incorporated into the base unit 106. Thescintillating fiber array coupled to a photodetector readout comprises astructured detector that can be deployed in place of an edge-on detectorin a CT scanner. Adjacent structured detectors such as this can bepositioned in a continuous, partially-offset or completely offsetconfiguration.

This ring detector geometry comprised of an array of 1Dscintillator-photodetector detectors oriented parallel to the axialdirection can be extended to multiple pixel widths along thecircumference, since planar or shaped entrance surface scintillatingfiber optic arrays and small, 2D high speed photodetector arrays areavailable. The use of 1D (or 2D) scintillator-photodetector detectorsmay offer advantages since manufacturing costs are typically reduced,although butting of 1D detectors is generally easier than butting of 2Ddetectors.

The same approach applies to a planar geometry concerning the use of 1Dor 2D scintillator-photodetector detectors. Although readout electronicssuch as ASICs can be attached to the 1D or 2D photodetector sensorsexternally, the readout electronics can alternatively be integrateddirectly on the substrate of the 1D or 2D photodetector sensors.

Orientation, Interaction Height and Sub-Aperture Resolution

Consider a scenario in which radiation is incident upon a planar edge-ondetector. The detector thickness (height) now defines the maximumdetector entrance aperture while the length or width of the detectorarea now defines the maximum attenuation distance for edge-on radiationdetector designs including semiconductor drift chamber, single-sidedstrip, and double sided strip detectors, including micro-strip detectorversions.

The interaction position along the height of the edge-on detectoraperture will be referred to as the interaction height. When ascintillator, semiconductor, gas, or liquid detector is irradiatedface-on the 1D positional information along the thickness direction ofthe detector is referred to as the interaction depth. Theelectronically-measured face-on detector DOI positional informationdefines the edge-on detector sub-aperture resolution (SAR).

Strip widths can be tapered or curved if focusing is desired. In thecase of double-sided parallel strip detectors in which opposing stripsare parallel, both electrons and holes can be collected to provide 2Dposition information across the aperture. If strips on one side runperpendicular to those on the other side, then depth-of-interactioninformation can be obtained. If strips are segmented in either asingle-sided or double-sided parallel strip detector thendepth-of-interaction information can be obtained and readout rates canbe improved.

In the case of double-sided parallel strip detectors (in which opposingstrips are parallel) or crossed strip or 2D pixelated array detectors,both electrons and holes produced by a radiation event can be collectedto provide 1D positional information between the anode and the cathodesides of the aperture. This 1D positional information is used todetermine electronically the sub-aperture spatial resolution.

Sub-aperture spatial resolution can be achieved by measuring either thetransit times of electrons and holes to the anodes and cathodes,respectively, or the ratio of anode and cathode signals. A significantbenefit may be gained by implementing sub-aperture resolution (e.g.,resulting from electronically-defined detector elements) because theedge-on detector aperture height no longer limits spatial resolutionalong that direction. Furthermore this 1D positional information may, insome situations, be used to estimate meaningful corrections to theexpected signal losses as a function of interaction height and thusimprove energy resolution. Other benefits include an increase inavailable image detector volume due to a decrease in the number ofedge-on detector physical boundaries (detector material propertiestypically degrade near the perimeter) and the number of gaps that may bepresent between edge-on detector planes.

The benefits of sub-aperture resolution (increased spatial resolution,signal loss compensation, fewer readout detectors, increased detectorvolume) that are possible with edge-on semiconductor detectors can alsobe attained using scintillator arrays in an edge-on detector geometry.Depth-of-interaction and interaction height information (e.g., forsub-aperture resolution) can be acquired using 1D and 2D scintillatorarrays, for example by adding dual-readout (photodetector readout)capability.

The semiconductor detector DOI accuracy is affected by parameters suchas the detector depth, electron and hole mobility, signal diffusion, andthe number of defects (such as traps) in the bulk semiconductormaterial. The specific parameters that affect scintillator detector DOIaccuracy vary with the DOI measurement technique.

Coupling a 2D photodetector readout array to the side of an edge-onscintillator array permits an analysis of the relative signal strengthmeasured at both ends of individual scintillator elements in the array.By calibrating the relative signal strength versus interaction locationin the direction of the aperture (interaction height), sub-apertureresolution can be achieved. With sufficiently fast readout detectors,time-of-flight measurements could also be used to determine theinteraction location. Thus, sub-aperture resolution can be attained for1D and 2D edge-on scintillator detectors, and a 2D, edge-on scintillatorarray detector can function as a 3D, edge-on scintillator arraydetector.

In many instances the one-side readout implementations of edge-on SARdesigns emulate the face-on DOI designs. In both face-on DOI and edge-onSAR scintillator detector designs, a one-sided or a multi-sided readoutcan be implemented. Thus, encoding techniques developed for one-sided ortwo-sided (or multi-sided), face-on DOI scintillator elements can beapplied to edge-on SAR scintillator elements. Furthermore, edge-on oredge-on with face-on 2D photodetectors coupled to two or more adjacentfaces of a scintillator block (e.g., a block geometry) can be employedto implement a 3D scintillator block detector using encoding techniques(e.g., providing SAR and DOI information).

A potential problem with face-on DOI scintillator detectors is thatCompton scattering of incident radiation is biased in the forwarddirection such that the probability of detecting the scatter eventdownstream from the initial off-axis event within the same scintillatormay not be small (resulting in an inaccurate DOI estimate). The edge-onSAR scintillator detector design reduces the likelihood that a Comptonscatter photon will be detected in the same scintillator for arelatively large range of incident angles. This simplifies the trackingof most subsequent interactions or events after a primary interaction.

The number of edge-on scintillator or semiconductor detector planesrequired to assemble an edge-on detector module can be reduced byimplementing the techniques developed for measuring the depth ofinteraction (DOI) within face-on scintillator and semiconductordetectors. The benefits of this approach can be illustrated byconsidering a scenario in which radiation is incident face-on upon theanode or cathode side of a planar semiconductor detector of known depthor thickness (height). The DOI spatial resolution can be determined bymeasuring either the transit times of electrons and holes to anodes andcathodes, respectively, or the ratio of anode and cathode signals.

Radiation incident approximately perpendicular to the plane orirradiation from the left or right side (approximately parallel to theplane) of an edge-on detector array is also allowed. Theside-irradiation geometry may be useful for specific applications. Forexample, it may be desirable to collimate the radiation so that thedetector region near the base and relevant readout electronics areremoved from direct irradiation. In addition, irradiation from the rightor left side would allow two edge-on detector arrays to be oriented suchthat one array faces the other array in close proximity. In general,spatial and energy resolution may be enhanced if sub-aperture heightinformation is acquired for edge-on detectors that are irradiated fromthe side.

FIG. 7 illustrates a perspective of a detector imaging system 1000 witha one-dimensional (1D) edge-on structured mold detector 700. As shown inFIG. 7, radiation 109 is incident onto the top surface of detector 700,in an edge-on 1D pixelated structured mold detector (silicon block)configuration. Holes 710 of structured mold detector 700 are filled withsemiconductor quantum dots or semiconductor detector materials. Channels720 are filled with attenuation material. These features are not toscale.

In this view, anode face 730 is oriented toward the front of detector700, showing three separate anode elements 735 separated by theattenuating material in channels 720. Cathode face 740 is orientedtoward the back of detector 700, with one or more cathode elements 745.Holes 710 and channels 720 can be etched or micromachined into (e.g.,silicon block) detector 700. For example, channels 720 may extend fromanode face 730 partially through detector 700 toward cathode face 740,as shown in FIG. 7, or channels 720 may extend completely throughdetector 700 to (or through) cathode face 740. Alternatively, the front(anode) and back (cathode) faces 730 and 740 can be reversed, and theplacement, arrangement, and configurations of channels 720 and holes 710may vary.

In the particular embodiment of FIG. 7, structured mold detector 700 hasholes 710 and channels 720 in which holes 710 are filled with quantumdot materials or other semiconductor detector materials. Channels 720are filled with an attenuating material to isolate or help isolateneighboring pixels 760. Detector 700 is irradiated in an edge-onconfiguration, with radiation 109 incident from the top, and front andback faces 730 and 740 of detector 700 are represented by conductiveanode and cathode elements 735 and 745, respectively. The top(front-end) and bottom (back-end) layers of detector 700 can also beinterchanged, without loss of generality.

For illustrative purposes, an implementation of detector 700 with only asingle layer of pixels 760 is shown. A cathode sheet covers back(cathode) face 740 of detector 700, and an anode sheet covers front(anode) face 730. Both the anode sheet on front face 730 and the cathodesheet on back face 740 can be segmented to create individual pixels 760in detector 700.

Selection of the radius (or other dimensions) of holes 710 and the depthand width of channels 720 is influenced by transport properties of theinformation carriers (or attenuation properties of the material if usedfor isolation purposes), and these quantities are selected in a suitablerange for detector 700 to identify interactions of radiation 109 withinpixels 760. In PbS quantum dots, for example, excitons have a range ofabout 20 nm, and holes 710 may be selected with a radius of about 50 nm,so that the excitons have acceptable probability of reaching thePbS-silicon heterojunction. Cost is also a design issue for choice ofquantum dot or semiconductor detectors materials (such as amorphous orpolycrystalline semiconductor materials).

FIG. 8 illustrates a perspective of a detector imaging system 1000 witha two-dimensional (2D), layered, edge-on structured mold detector(silicon block) 700. In this configuration, channels 720A are providedwith relatively low-Z detector material, e.g., in the first or top layer701 of detector 700, and channels 720B are filled or provided withrelatively moderate or high-Z detector material, e.g., in the second orlower layer 702 of detector 700. These features are not to scale.

Channel 720C is provided with a filter material, e.g., between top(low-Z) layer 701 and bottom (moderate or high-Z) layer 702 of detector700. Anode elements 735 are segmented both by layer 710, 702 and bypixel 706 within each layer 701, 702. Cathode elements 745 can besimilarly divided.

In the particular embodiment of FIG. 8, structured mold detector 700 haschannels 720A and 720B filled with semiconductor quantum dots orsemiconductor detector materials. Hole structures are not necessarilyrequired. As shown, two layers 701, 702 of pixels 706 are provided, withrelatively low-Z (or lower-Z) detector materials in channels 720A offirst layer 701 (e.g., the top layer, where radiation 109 is incidentonto edge-on detector 700), and moderate or relatively high-Z (orhigher-Z) materials in second layer 702 (e.g., the bottom layer ofdetector 700, reached by radiation 109 passing though top layer 701).

For illustrative purposes, the implementations of anode face or layer730 and cathode face or layer 740 are represented with only one level ofpixels 706 per detector layer 701, 702. Alternatively, there may bemultiple pixels 706 per detector layer, and/or multiple anode andcathode elements 735 and 745. The selection of channel width (or holeradius) for detector system 700 is influenced by the transportproperties of the information carriers, as described above.

FIG. 9 illustrates a perspective of a multilayer detector system 800 fora CT and/or PET detector imaging system 1000. In this particularexample, detector system 800 includes N=4 (four) individual layers: atop or front-end layer 801, two middle layers 802A and 802B, and abottom or back-end layer 803.

As shown in FIG. 9, the first (top or front-end) layer 801 is formedwith an array of edge-on, 2D pixelated detectors, e.g., using arelatively low-Z semiconductor detector material such as silicon. Thetwo middle layers 802A and 802B are formed with arrays of face-on 1D or2D pixelated detectors, e.g., using a moderate-Z semiconductor detectormaterial such as CZT or CdTe. The bottom (or back-end) layer 803 isformed with an array of edge-on, 2D pixelated detectors, e.g., using amoderate or relatively high-Z scintillator or semiconductor detectormaterial.

The dimensions of pixels within a layer or within different layers maybe different. Therefore, the spatial resolution (as well as temporal andenergy resolution) properties of individual detector layers (as well asthe pixels within a detector layer) need not be the same and aredictated by the imaging requirements as well as cost. This principleapplies for one detector layer, two detector layers, three detectorlayers, four detector layers, etc. For example, although detector layer803 pixels are depicted with the same surface area (e.g., 1×1) as pixelsin detector layers 801 and 802, imaging and cost requirements mayindicate (or dictate) a different relationship A×B where each of A and Bcan be less than, equal to or greater than one (including blockdetectors).

Depending on application, first (top) layer 801 can be used for CT andPET imaging (e.g., employing a combination of x-ray and Compton scatterinteractions). Middle layers 802A and 802B could also be used for CTand/or PET imaging, while bottom layer 803 can be used primarily for PETimaging. In some embodiments, middle layers 802A and 802B are providedas removable/insertable units, which are configured for insertion intoand removal from imaging system 1000 between top and bottom layers 801and 803 of detector 800. Bottom (PET) layer 803 can also be provided inthe form of a 3D sub-aperture resolution (SAR) detector, for example asdescribed in Nelson, U.S. Pat. No. 7,635,848. In general, the N detectorlayer design is appropriate for at least one of Compton (including gammacamera/SPECT) imaging, Compton-PET imaging, PET (including TOF PET)imaging, high resolution CT imaging, CT imaging. Either energyintegration or PC or PCE capability can be implemented in individualdetector layers depending on the detector capabilities and imagingrequirements.

FIG. 10A illustrates a perspective of an alternate multilayer detectorsystem 800 for, in one implementation, a CT and/or PET detector imagingsystem 1000. In this particular example, detector system 800 includesN=3 (three) layers: first (top) layer 801, second (middle) layer 802 andthird (bottom) layer 803.

As shown in FIG. 10A, first (top or front-end) layer 801 is formed withan array of edge-on, 2D pixelated semiconductor detectors, e.g., using arelatively low-Z semiconductor detector material such as silicon. Second(middle) layer 802 is formed with an array of edge-on pixelateddetectors, e.g., using a moderate-Z semiconductor or scintillatordetector material (alternatively, layer 802 can be implemented using anarray of face-on pixelated detectors). Third (bottom or back-end) layer803 is formed with an array of edge-on, 2D pixelated detectors, e.g.,using a moderate or high-Z detector material. Second (middle) detectorlayer 802 can also be configured in removable/replaceable form (e.g., ifdetector layer 803 is not implemented then N=2 layers), and third(bottom) layer can be replaced (for example) by a 3D DOI detector or a3D SAR detector, as described above for the four-layer embodiment ofFIG. 9. Depending upon embodiment, one goal is to provide a three-layerconfiguration including an edge-on Si, face-on scintillator orsemiconductor detector layer and an edge-on semiconductor orscintillator detector layer (e.g., in the front-end and/or back-endlayers). A face-on middle detector layer could include an integrator tohandle high fluence, e.g., in an x-ray CT imaging system (or combinedCT/nuclear medicine system), with the awareness that 511 keV photons maynot be entirely contained. Thus, such a face-on layer may providelimited energy resolution, or incorporate photon counting or integrationcapability, as described herein.

FIG. 10B illustrates a perspective of CT and/or PET detector imagingsystem 1000 with a face-on back-end detector layer 803. As shown in FIG.10B, first (top or front-end) layer 801 and one or more middle layers802 of detector system 800 are formed with arrays of edge-on, 1D or 2Dpixelated detectors. Third (bottom) layer 803 is formed with an array offace-on, 1D or 2D pixelated detectors.

FIG. 10C illustrates a perspective of a multilayer CT and/or PETdetector imaging system 1000 with a face-on back-end detector layer.FIGS. 10B and 10C show one or more middle detector layers 802 ofdetector system 800 are formed with arrays of edge-on, 1D or 2Dpixelated detectors.

FIG. 10D illustrates a perspective of a multilayer CT and/or PETdetector imaging system 1000 with a face-on back-end detector layer.FIG. 10D illustrates a middle detector layer 802 of detector system 800formed with an array of face-on pixelated detectors.

FIGS. 10B, 10C and 10D illustrate CT and/or PET detector imaging system1000 with an edge-on first (top or front-end) detector layer 801, one ormore middle detector layers 802 and a face-on back-end detector layer803. As shown in FIGS. 10B, 10C and 10D, first (top or front-end)detector layer 801 of detector system 800 is formed with arrays ofedge-on, 1D or 2D pixelated detectors. FIGS. 10B and 10C show one ormore middle detector layers 802 of detector system 800 formed witharrays of edge-on, 1D or 2D pixelated detectors. FIG. 10D illustratesmiddle detector layer 802 of detector system 800 formed with an array offace-on pixelated detectors.

FIGS. 10B, 10C and 10D show a third (bottom) detector layer 803 formedwith face-on block, 1D, 2D (or 3D if DOI is implemented) pixelateddetectors. As described herein, one or more of the layers could combineedge-on and face-on detector elements (varying at least one of spatialresolution, energy resolution, temporal resolution and stopping powerwithin a detector layer). The dimensions of pixels within a layer orwithin different layers may be different. The spatial resolution (aswell as temporal and energy resolution) properties of individualdetector layers (as well as the pixels within a detector layer) need notbe the same and are indicated (or dictated) by imaging requirements aswell as cost.

This principle applies for one detector layer, two detector layers,three detector layers, four detector layers, etc. For example, in oneimplementation the surface area of a pixel of third detector layer 803as shown in FIG. 10C is depicted, for illustrative purposes, as being(approximately) the surface area of a 5×5 array of pixels in firstdetector layer 801 or middle detector layer 802. In this instance thepixel size employed in detector layers 801 and 802 may be appropriatefor high resolution CT imaging (with a pixel surface dimension rangingfrom approximately 0.2-1.0 mm at this time) whereas the pixel sizeemployed in detector layer 803 may be appropriate for at least one ofCompton (including gamma camera/SPECT) imaging, Compton-PET imaging, PET(including TOF PET) imaging, CT imaging. Other implementations mayresult in third detector layer 803 pixels being smaller, the same orlarger than pixels in first detector layer 801 pixels or second detectorlayer 802.

FIG. 10D depicts detector layer 803 with a pixel area which is moreblock-like. In this instance detector layer 803 might be comprised ofone or more blocks used primarily for timing and/or energy determinationfor scattered PET photons from detector layer 801 or detector layers 801and 802 (if detector layer 802 is not present then this represents animplementation of an N=2 detector layer imaging system 1000). (Note thatvarious implementations of block detectors are possible including simple1D block detectors, 2D block detectors (including, but not limited to,1D arrays or gamma cameras) and 3D block detectors). Features includingfast or very fast temporal response, present in one or both layers ofthe above-described two layer, coincidence (including TOF) PET detectorsystem implementation, can be present in one or more of the detectorlayers for the three-layer CT and/or PET, Compton-PET, and Compton(including gamma camera) detector imaging systems. Either energyintegration or PC or PCE capability can be implemented in individualdetector layers depending on detector capabilities and imagingrequirements.

Middle detector layer (or layers) 802 may be provided in a removableconfiguration and bottom (back-end) detector layer 803 may be replacedwith an SAR detector, as described above. In addition, the orientationof first (front-end) and third (back-end) layers 801 and 803 can beinterchanged with respect to the direction of incident radiation 109,without loss of generality.

FIG. 11 illustrates a perspective of a detector imaging system 1000 witha focused two-dimensional (2D), layered, edge-on, pixelated structuredmold (silicon block) detector 700. As shown in FIG. 11, individualpixels 706 diverge in width along the direction of incident radiation109, e.g. with respect to the direction of radiation 109 from adiverging source such as an internal radionuclide or a diverging x-raybeam. This provides pixels 706 and detector 700 with a focused structuregeometry, as described herein.

Channels 720A are provided with relatively low-Z detector material,e.g., in the first or top layer 701 of focused detector 700, andchannels 720B are filled or provided with relatively moderate or high-Zdetector material, e.g., in the second or lower layer 702 of focuseddetector 700. Channel 720C is filled or provided with a filter material,e.g., between top (low-Z) layer 701 and bottom (moderate or high-Z)layer 702 of focused detector 700. Anode elements 735 and cathodeelements 745 can be segmented both by layer 710, 702 and by pixel 706within each layer 701, 702.

In the particular embodiment of FIG. 11, focused, structured molddetector 700 has channels 720A and 720B filled with semiconductorquantum dot materials or semiconductor detector materials. Holestructures are not necessarily required. As shown, two layers 701, 702of pixels 706 are provided, with relatively low-Z materials in channels720A of first layer 701, and moderate or relatively high-Z materials insecond layer 702. Anode layer 730 and cathode layer 740 may havemultiple anode and cathode elements 735 and 745 per detector layer, withmultiple pixels in one or both of layers 701 and 702.

Discrete and Semi-Continuous Structured 3D Scintillator Detectors

Nelson (U.S. Pat. No. 7,635,848) teaches a planar, discrete structured3D scintillator detector modules (that can be stacked) using crossed topand bottom (planar) layers of parallel arrays of discrete scintillatorrods with light sharing and an optical readout at one end of each arrayof discrete scintillator rods. Various photodetector implementations(photodiodes and amplified photodetectors including, but not limited to:PMTs, PSPMTs, APDs, SiPMs, silicon nanowires, iDADs, Se-APDs, GaAsPMs,DiamondPMs, microchannel plates, etc.) and electronic readoutimplementations (integrating, photon counting, photon counting withenergy resolution, coincidence) have been described.

Photodetector configurations can include, but are not limited to,discrete pixel photodetectors, photodetector strips with readoutcircuitry at the ends of the strips (measuring relative signal strengthand/or propagation time), photo-emissive detector strips and areaphotodetectors. Strip and area photodetectors include, but are notlimited to, pixelated PMTs (PSPMTs) photodetectors, 1D/2D pixelated,strip or area APD/SiPM/iDADs/silicon nanowires/microchannel platephotodetectors (as well as implementations utilizing Se-APDs, GaAsPMs,DiamondPMs, etc.). Area photodetectors can implement readout circuitryat the four corners.

Variations on this discrete structured 3D scintillator detector moduledesign implement two (or more) layers. A two layer detector module canemploy top and bottom layers of the same or different scintillator rodmaterials, discrete scintillator rods in the top and bottom layers withthe same or different spatial dimensions, shared photodetector sensors,applied reflective materials and patterns, applied absorptive and/orscattering materials and patterns, applied wavelength shifting (WLS)materials (including conventional fluorescent dopants as well as quantumdots and other nano-particles and microdots) and patterns (including WLSmaterials combined with reflectors), WLS fibers (including WLS fiberscombined with reflectors), structured surfaces, etc. (see Nelson, U.S.Pat. No. 8,017,906). This approach can be extended such thatscintillator rod parameters (material, physical properties, dimensions)within a layer and between layers can be varied. Note thatimplementations of quantum dots, other nano-particles and microdots canfunction as wavelength shifting materials and/or direct ionizingradiation conversion materials and can be employed on the surface aswell as within the volume of a scintillator pixel, rod, sheet, block orfiber.

Thus, different scintillator materials can be employed as well asvariations of a scintillator material (different properties such astemporal decay characteristics, spectral distribution, conversionefficiency, light transport directionality and/or absorption, etc.).Furthermore, arrays of discrete scintillator pixels can replace one ormore discrete rods (or all rods) within a layer. Scintillator pixelparameters (spatial dimensions, scintillator materials, physical andvirtual internal and external structures, the application of WLSmaterials and patterns of WLS materials, etc.) within a layer andbetween layers can be varied.

One or more layers of discrete scintillator rods and/or discretescintillator pixels can be replaced by a scintillator sheet or multiplescintillator sheets which can be continuous or structured (physical andvirtual structures can be introduced into, or onto, scintillatorsheets). Physical and virtual internal structures include, but are notlimited to, rods and/or pixels. Virtual internal structures can include,but are not limited to, vertical barriers, horizontal barriers, angledbarriers, grids, fibers and dispersive formations. Physical and virtualsurface and/or internal structures can be formed by roughening, sawing,drilling, sub-surface laser engraving, ion implanting, photolithography,deposition, gluing, embedding, etc. Scintillator sheet parameters withina layer(s) or between layers can be varied. A single scintillator sheet(continuous or structured) can be positioned between two layers ofdiscrete scintillator rods and/or pixels. Physical and virtual surfaceand/or internal structures can be employed with scintillator sheets,scintillator rods and any coatings.

Multiple scintillator sheets can be coupled to implement a 3Dscintillator detector module. A 3D scintillator detector module stackcan incorporate 3D scintillator detector modules with the same ordifferent properties. WLS materials and (encoded) patterns of WLSmaterials can be applied to one or more scintillator rod, pixel, sheetsurfaces. Scintillator sheets, as well as layers of discretescintillator rods and/or pixels can implement at least onenon-scintillator supporting layer to provide at least one of: additionalstructure integrity, reflection and/or redirection of the fluorescencesignal, WLS materials (including encoded patterns of WLS materials),radiation filtering/conversion.

Physical or virtual structures can be introduced into, or onto, one ormore discrete scintillator rods (as well as discrete scintillator pixelsand scintillator sheets, if present) in order to improve spatial and/ortemporal and/or energy resolution. Top and bottom layer scintillator rod(or pixel) surfaces that contact each other (as well as readout surfacesand other readout surfaces) can be modified to increase and/or directlight transmission by implementing at least one of: physical or virtualsurface microstructures and/or macrostructure patterns attached to,deposited onto or cut/etched into or implanted (by ion implantation,sub-surface laser engraving or other techniques) into the scintillatorcrystal surface(s), surface macrostructures patterns, surface coatings,WLS materials including, but not limited to, WLS quantum dots and WLSfilms (see Nelson, U.S. Pat. No. 8,017,906). For example, rough surfacepatterns can be employed to diffuse directional light transmissionwhereas more sophisticated patterns such chevron designs (curved orangled channels) can be employed to help direct light transmissionbetween crossed scintillator rods. Cherenkov radiation can also berecorded. Various techniques have been described for creating structureswithin or on the surface of scintillator rods (and therefor scintillatorpixels), rod coatings, etc., and these techniques are readily extendedto structured and continuous scintillator sheets and blocks.

Another variation of the discrete structured 3D scintillator detectormodule design (applicable to the other 3D detector designs describedherein) implements discrete scintillator rods (or scintillator pixels orscintillator rods and pixels) or a continuous or structured (withdiscrete or virtual rods and/or pixels) scintillator sheet in which oneor more scintillator properties of a scintillator rod (or scintillatorpixels or a scintillator sheet) such as temporal decay distribution,spectral distribution, conversion efficiency is varied spatially alongthe length of a scintillator rod (or between scintillator pixels or overthe area of a scintillator sheet) according to an encoded pattern.Positional information can thereby be encoded into the scintillator rods(or scintillator pixels or a scintillator sheet). The implementation ofencoded scintillator rods (or scintillator pixels or scintillator sheet)in structured 3D scintillator detectors may be used to enhanceresolution for crossed (or uncrossed) scintillator rod layers, orscintillator pixels layers or (continuous or structured) scintillatorsheets.

Yet another method of encoding position information along the length ofa scintillator rod for the discrete structured 3D scintillator detectormodule is to apply continuous or discrete patterns (including 1D and 2Dpatterns) of WLS materials along the lengths of the rods such thatwavelength and/or pulse shape properties (including time delays inemission) vary with position (see Nelson, U.S. Pat. No. 8,017,906). Thisencoding method can be implemented with scintillator pixels. Two or moreWLS materials can be present at a location within the pattern. Thisdesign can be extended to the case of fibers crossed with scintillatorrods (or pixels) in which scintillator rods (or pixels) are encoded,fibers are encoded or fibers and scintillator rods (or pixels) areencoded. Furthermore, this multiple-WLS design can be combined with (orreplaced by) one or more of absorbers, scatterers, reflectors, surfaceand/or internal implementations of physical and/or virtual structures(including implementations employing patterns) and can be employed witha single layer of rods with one sided or two-side readout to implement asingle layer 2D or 3D scintillator detector that can be employed in asingle layer or multiple layer detector module. Patterns can beimplemented internally and/or on at least one side of individual rods ina single layer. An adjacent layer can be optically coupled with oroptically isolated from this single layer of rods. For example, if theadjacent layer is continuous or pixelated then optical coupling istypically employed (the adjacent layer can be an intermediate layer). Ifthe adjacent layer is comprised of rods parallel to the single layerrods then optical isolation is typically employed. If the adjacent layeris comprised of rods at an angle (such as perpendicular) to the singlelayer rods then optical isolation or optical coupling (if beneficial)can be employed.

Furthermore, this multiple-WLS design (and variants thereof) can be usedwith scintillator slabs and blocks as well as scintillator fibers.Potential benefits include improved spatial resolution, energyresolution and/or reduced ambiguity concerning the detection of multipleevents within the coincidence resolution limits of the discretestructured 3D scintillator detector(s). For example, a fiber array canbe shared across multiple discrete structured 3D scintillator detectors.This approach to encoding can be readily extended to a continuous orstructured scintillator sheet (including a rod and/or pixelatedscintillator sheet) by implementing a 2D pattern of WLS materials.Multiple-WLS designs, in general, can be combined with (or replaced by)one or more of absorbers, scatterers, reflectors, surface and/orinternal implementations of physical and/or virtual structures(including implementations employing patterns). Physical internalstructures can include embedded quantum dots, nano-particles andmicrodots.

Furthermore, it can be employed with other discrete or shared-lightoutput 2D and 3D detector designs (including thick scintillator sheetdetectors) implemented for nuclear medicine and/or PET imaging. Spatialpositioning can be determined by employing photodetectors at both endsof an array of scintillator rods within a layer (or a scintillatorsheet) using signal weighting to compare signal strength at both endsand/or TOF resolution techniques to compare pulses arrival times at bothends. Spatial positioning can also be determined by employingphotodetectors at one end of an array of scintillator rods within alayer (or a scintillator sheet) using TOF resolution to compare arrivaltimes of the initial pulse with the direct reflected pulse or aneffective reflected pulse (a delayed signal) from the opposite end.(Note that an effective reflected pulse (delayed signal) can be providedby a (passive) WLS material positioned at the scintillator face oppositethe photodetector readout face is but one technique for providing adelayed signal. Other passive or active photo-activated signal delaymechanisms can be implemented.)

Sources of scintillator rods include a discrete 1D array of scintillatorrods, (one multiple layer implementation could employ a discrete 2Darray of scintillator rods), a structured scintillator sheet, asingle-sided semi-continuous structured scintillator sheet, adouble-sided semi-continuous structured scintillator sheet, adouble-sided semi-continuous structured scintillator sheet with anintermediate layer. In addition, this approach can be readilyimplemented with scintillator fiber layers (or 2D arrays) as well asfiber optic layers (or 2D arrays). One or more non-scintillatorsupporting layers (as previously described herein) can be employed.

Patterns of light sharing between readout photodetectors may be used toimprove estimates of depth of interaction (DOI) or sub-apertureresolution (SAR). One or more encoding techniques can be implementedwith the detectors taught herein. A previously herein-describedvariation on the discrete structured 3D scintillator detector module isa discrete rod-structured, pixel-structured 3D scintillator detectormodule which implements a parallel array of discrete scintillator rodswith light sharing and optical readouts at one or both ends of the arrayof discrete scintillator rods coupled to a second layer comprised of adiscrete 2D pixel scintillator array. (Note that implementations caninclude not only arrays of rods or 2D arrays of pixels but also mixedarrays of rods and pixels.) Furthermore, a discrete pixel-structured,pixel-structured 3D scintillator detector module can be implemented inwhich both layers employ 2D arrays of discrete scintillator pixels (thetwo layers of discrete 2D scintillator pixel arrays can be aligned oroffset). One or more encoding techniques can be implemented.

The discrete 2D pixel scintillator array layer can be replaced with ascintillator sheet (a structured scintillator sheet) with an imposedpixel structure including at least one of physical pixels(cut/sawed/etched) and virtual scintillator pixels (for example, via ionimplantation). The physical and virtual pixels need not extend throughthe entire thickness of the scintillator sheet layer. Variations in thephysical or virtual gap (divider) depth used to define a scintillatorpixel may be employed in order to modify (locally) the distribution offluorescence (or other optical) signals. Furthermore, encodingtechniques such as spatially varying scintillator properties and/orapplying continuous or discrete patterns (including 1D and 2D patterns)of WLS materials along the lengths of the rods or on the pixels of the2D pixel arrays such that wavelength and/or pulse shape properties(including time delays in emission) vary with position can be used toenhance resolution.

Additional variations on the use of a scintillator sheet layer include ascintillator sheet layer with an imposed rod array structure (replacinga layer of discrete scintillator rods in a crossed rod geometry or in adiscrete rod-structured, 2D pixel-structured geometry), or ascintillator sheet layer with no structure. For the case of ascintillator sheet layer with an imposed rod structure an opticalreadout can be positioned at one end of each array of scintillator rodsor the optical readouts can be positioned at both ends of the array ofdiscrete scintillator rods. For the case of a scintillator sheet layerwith no structure the optical readouts can be positioned at one end orboth ends of the array of discrete scintillator rods.

The implementation of spatial encoding through techniques such asspatially varying scintillator properties and/or applying continuous ordiscrete patterns of WLS materials on one or more surfaces for discretescintillator rods (and/or pixels) and a scintillator sheet layer(imposed rod/pixel structure, no structure) can be used to improveresolution and/or enable the use of optical readouts at a single end ofthe array of discrete scintillator rods. It should be understood thatencoding techniques described herein can be employed with any of theplanar or non-planar scintillator configurations (discrete scintillatorrod, discrete scintillator pixel array, structured or continuousscintillator sheet, scintillator block, fiber arrays, scintillator fiberarrays, etc.) taught in this specification.

Photodetector configurations can include, but are not limited to,discrete pixel photodetectors, photodetector strips with readoutcircuitry at the ends of the strips (measuring relative signal strengthand/or propagation time), photo-emissive detector strips and areaphotodetectors. Strip and area photodetectors include, but are notlimited to, pixelated PMTs (PSPMTs) photodetectors, 1D/2D pixelated,strip or area APD/SiPM/iDADs/silicon nanowires/microchannel platephotodetectors (as well as implementations utilizing Se-APDs, GaAsPMs,DiamondPMs, etc.). Area photodetectors can implement readout circuitryat the four corners. Note that applications include not only medicalx-ray and gamma ray detection and imaging but also the detection andimaging of charged and uncharged particles in general.

FIG. 12A shows a discrete structured 3D scintillator detector module 201with crossed top (first) and bottom (second) layers that implementdifferent scintillator rod materials with different dimensions 202, 203(non-uniform discrete structured 3D scintillator detector module). (Analternative implementation uses different scintillator materials butwith the same dimensions.) Furthermore, the scintillator rods 202, 203within a layer may implement scintillator rod materials and/ordimensions with the same or different properties. Scintillator rodlengths in the top and bottom layers can be the same or different.Physical structures or virtual structures (using ion implantation,subsurface laser, etc. techniques) that modify optical propagation canbe introduced into, onto and between scintillator rods (within a layerand between layers) in order to improve at least one of event positionalaccuracy, timing and detected signal levels. The physical structures(linear and non-linear) can include, but are not limited to, airstructures, coating structures, coupling material structures, photoniccrystal structures, nano-layered metamaterials (including nanocavities)structures, refracting structures, diffracting structures, lensstructures, micro-spheres and related micro-structures, fiber plates,WLS material structures, specular reflector structures, diffusereflector structures, absorbing structures, scattering structures,polished surfaces, roughened surfaces, etched surfaces, shaped surfaces.

The detector module 201 could be used to detect and/or track photonsand/or particles with different energies (different properties),including scattered photons and/or particles. 3D spatial resolution isdifferent for events detected in the top and bottom layers. Both uniformand non-uniform discrete structured 3D scintillator detector modules 201can be positioned (stacked), one on top of another, to form a stackeddiscrete structured 3D scintillator detector module. Individual rodsurfaces can be polished or unpolished (or a patterned combination ofpolished and unpolished). One or more of the non-readout surfaces of therods (or segments of rod surfaces) can be coated and/oretched/roughened, notched, etc. in order to direct more light to thepreferred readout photodetector(s). Zero, one or multiple scintillatorcoatings and/or surface treatments (including surface treatments to thescintillator coatings) or combinations thereof can be applied incontinuous or discrete patterns to the individual rod surfaces.

Optionally, patterns of reflective and/or absorptive optical barrierscan be introduced within one or more of the scintillator coatings tocontrol the propagation of the scintillator and/or WLS optical photonswithin the scintillator coating or coatings and between scintillatorlayers. A simple optical barrier pattern could be a grid patternimplemented between 2 layers of crossed rods with the same or differentrod cross sections. The optical barrier grid pattern corresponds to thepattern formed by the crossed rods (for example, crossed rods with 1×1mm² and 2×2 mm² cross sections would implement an optical barrier gridpattern with element dimensions of 1×2 mm² or finer).

Scintillator coatings can include, but are not limited to, one or morelow index coatings, high index coatings, reflective coatings, absorptivecoatings, diffuse coatings, diffuse reflective coatings, WLS coatings.Coatings can be applied uniformly or non-uniformly, including continuousand discrete patterns of coatings. Physical and/or virtual internalstructures can be introduced into or onto coatings and/or scintillatorrods by roughening, sawing, drilling, sub-surface laser engraving, ionimplanting, photolithography, deposition, gluing, embedding, etc. tohelp direct (and/or encode) the distribution of the fluorescence signalto the preferred readout photodetector(s). For example, an embeddedphysical internal structure could include a discrete or continuouspattern of quantum dots (and/or other suitable nano-particles ormicrodots) distributed internally along the length of a scintillator rodsuch that one or more properties (such as emission spectrum and/ortemporal response) of the quantum dots and/or density vary with 2D (or3D) position within the scintillator rod (a technique readily extendedto encoding of scintillator sheets and blocks as well as scintillatorfibers). Examples of quantum dot patterns include, but are not limitedto, varying quantum dot density versus position (a gradient) along atleast one dimension (such as the length) of the scintillator rod,varying quantum dot property versus position along at least onedimension (such as length) of the scintillator rod, varying quantum dotdensity and property versus position along at least one dimension (suchas length) of the scintillator rod. Embedded physical internalstructures can also be implemented within scintillator pixels, sheets,blocks and fibers. Non-imaging optics techniques, photonic crystalstructures, nano-layered materials (including nanocavities) structures,WLS structures, refracting structures, diffracting structures, lensstructures, micro-sphere structures, etc. can be applied.

For example, in one implementation a relatively high index scintillatorrod can be coated on one or more surfaces with at least one relativelylow index of refraction coating in order to exploit high reflectivityfor angles of incidence near or greater than the critical angle (thefiber optic effect). A WLS coating or WLS structures, in one instance,can be applied to the rod surface in contact with the other scintillatorrod layer to modify (in this case broaden) the angular distribution ofthe incident light and shift the fluorescence spectrum. Other techniqueof modifying the angular distribution include, but are not limited to,roughening the coating surface, creating/applying photonic crystalstructures, nano-layered metamaterials (including nanocavities)structures, refracting structures, diffracting structures, or lensstructures at the coating surface, applying microspheres or fiber platesat the coating surface. Coatings and/or physical (and/or virtual)structures can be applied to other rod surfaces to enhance at least oneof spatial resolution, energy resolution, temporal resolution.

These 3D scintillator detector designs (as well as other opticalradiation detector designs described herein) can be used to detectCherenkov and/or scintillation optical radiation. Alternativeimplementations employ relatively high-Z, non-scintillator materials todetect Cherenkov optical radiation.

FIG. 12B illustrates a perspective of a uniform discrete structured 3Dscintillator detector module 201 with crossed top and bottom layers thatimplement the same scintillator rod materials with the same dimensionswith the top (first) layer and bottom layer overhanging their commonsurface employed for light sharing. While the sides of the discretestructured 3D scintillator detector module 201 in FIG. 12A are shown asuniform, an alternative arrangement in FIG. 12B illustrates a uniformdiscrete structured 3D scintillator detector module with crossed top andbottom layers that implement the same scintillator rod materials withthe same dimensions 205 wherein the top layer and the bottom layers ofrods 205 extend (overhang) their common surface that is employed forlight sharing. These rod extensions can be very small (permitting thephotodetector to be offset from the other layer) or sufficiently largeso as to gain one or more extra pixels (defined by an absence of directlight sharing) while also providing an offset for the photodetector fromthe other layer.

One or more of the non-readout surfaces of the rod extensions can beprovided with a coating and/or etched, etc. in order to direct morelight to the readout photodetector. The implementation of scintillatorlayers that overhang their common surface is not limited to the use of asingle scintillator material or a single rod dimension, a double-sidedsemi-continuous structured scintillator sheet with a rod structure canimplement rod extensions. Furthermore, this approach is readily extendedto include scintillator sheet extensions.

FIG. 12C illustrates a discrete rod-structured, pixel-structured 3Dscintillator detector module 201 which implements a parallel array ofdiscrete scintillator rods 207 with light sharing and optical readoutsat one or both ends of the array of discrete scintillator rods in thebottom (second) layer are coupled to a top (first) layer comprised of anarray of discrete (or virtual) scintillator pixels 210.

The first and second (or top and bottom) layers can be interchangedwithout loss of generality. The scintillator materials employed in thefirst and second layer can be the same or different in terms of at leastone of fluorescence decay properties, fluorescence efficiency, radiationinteraction cross sections. In general, one or more additional(structured or continuous) scintillator layers can be inserted betweenthe top and bottom scintillator layers.

An alternative to implementing crossed layers (referred to as the topand bottom layers for convenience) of discrete scintillator rods is toimplement a version of a structured scintillator sheet by introducingphysical or virtual structures (such as scintillator rods) to define thetop (first) and bottom (second) layers of a scintillator sheet so thatthe top and bottom layers of scintillator rods defined within thescintillator sheet are crossed. This structured scintillator sheetrepresents one implementation of a semi-continuous structuredscintillator sheet. Multiple variations of a semi-continuous structuredscintillator sheet can be implemented including two layers and twolayers with an intermediate layer (which may be continuous or havestructures). Physical and virtual structures can be introduced into, oronto, semi-continuous structured scintillator sheets.

The top and bottom layers of a scintillator sheet are not limited toimplementing arrays of rods. For example, the bottom layer can implementrods whereas the top (first) layer can implement a 2D pixel array or becontinuous (no imposed structure). A bottom layer can implement a 2Darray of pixels whereas the top layer can be continuous (no imposedstructure) or a 2D pixel array (in this case the top and bottom layersof pixel arrays can be aligned or offset). Both rods and pixels can beimplemented within a top and/or bottom layer.

The dimensions of rod or pixels within an array can be uniform but neednot be uniform. Encoding techniques can include, but are not limited to,implementing scintillator in which one or more scintillator propertiesvary spatially and/or applying surface treatments such as continuous andpatterned WLS materials. A semi-continuous structured scintillator sheet(a single scintillator layer) can be employed directly as a 3Dscintillator detector module (since it is comprised of at least twosub-layers that can be modeled as scintillator layers). Asemi-continuous structured scintillator sheet can be employed with alayer of discrete scintillator rods or between two layers ofscintillator rods to form a 3D scintillator detector module. Asemi-continuous structured scintillator sheet can be employed with oneor more of: a scintillator sheet, a second continuous structuredscintillator sheet, a layer with discrete scintillator rods or a layerwith a discrete 2D pixel scintillator array to form a 3D scintillatordetector module.

The flexibility of this design enables multiple implementations thatpermit simplifications in manufacturing of this design compared to thediscrete structured 3D scintillator detector module design (althoughdesign attributes such as the ability to employ different scintillatormaterials within a layer or between layers require using scintillatorsheets with non-uniform properties). Cherenkov radiation can also berecorded. Increased light sharing between readout photodetectorscompared to typical discrete structured 3D scintillator detector moduleimplementations can be used to further improve estimates of depth ofinteraction (DOI) or sub-aperture resolution (SAR).

One implementation of a semi-continuous structured 3D scintillatordetector module is the double-sided, semi-continuous structured 3Dscintillator detector module that implements a single scintillator sheetinto which physical gaps (cut/sawed/etched) and/or virtual gaps (forexample, via ion implantation or sub-surface laser engraving) are madewith respect to the top and bottom surfaces creating semi-continuousdiscrete or virtual scintillator rods, respectively (a double-sidedsemi-continuous structured scintillator sheet with a rod structure).Since the depth of the rod structures of the top and bottom surfaces canbe controlled at least one of overlap (creating pixels in the overlapregion), contact or a gap between top and bottom surface rods can beimplemented in a desired pattern within the scintillator sheet. (Notethat a physical or virtual partial pixel structure can optionally beimplemented in a desired pattern within discreet scintillator rods.)

Arrays of semi-continuous discrete or virtual scintillator rods (orcombinations of semi-continuous and discrete scintillator rods) arethereby created within the top and/or bottom layers of the scintillatorsheet in order to form a double-sided, semi-continuous structured 3Dscintillator detector using a double-sided semi-continuous structuredscintillator sheet with a rod structure. As previously described herein,physical and/or virtual structures can be introduced into, or appliedto, one or more of the semi-continuous scintillator rods in order toimprove spatial and/or temporal and/or energy resolution. The crossingangle of the semi-continuous discrete or virtual scintillator rods istypically 90 degrees although other crossing angles can be implementedas needed.

Readout photodetectors are coupled to one end of each array ofsemi-continuous discrete and/or virtual scintillator rods comprising thetop and bottom layers. The dimensions (thicknesses, width) of thesemi-continuous discrete and/or virtual rods within the top and bottomlayers can be different (as is the case for the discrete structured 3Dscintillator detector) with the total thickness of rods from both layersbeing comparable to the thickness of the scintillator sheet.Furthermore, the width of individual rods within the top layer and/orbottom layer can be uniform or non-uniform.

The double-sided, semi-continuous structured 3D scintillator detectormodule can, for suitable applications, offer reduced manufacturing costscompared to the discrete structured 3D scintillator detector. Forexample, as a discrete scintillator rod cross section becomes finermanufacturing yields tend to decrease and assembly costs tend toincrease. As previously described herein, encoding techniques such asspatially varying scintillator properties and/or applying continuous ordiscrete patterns (including 1D and 2D patterns) of WLS materials, etc.can be used to enhance resolution.

Another variation on the double-sided, semi-continuous structured 3Dscintillator detector module is a double-sided, semi-continuousstructured 3D scintillator detector module with a rod structure and apixel structure which implements a single scintillator sheet into whichphysical gaps (cut/sawed/etched) and/or virtual gaps (for example, viaion implantation or sub-surface laser engraving) are made with respectto the top and bottom surfaces creating semi-continuous discrete orvirtual scintillator rods in one layer and semi-continuous discrete orvirtual scintillator 2D array of pixels in the other layer (adouble-sided semi-continuous structured scintillator sheet with a rodstructure and a pixel structure). In this implementation the readoutphotodetectors are coupled to one or both ends of the array ofsemi-continuous discrete and/or virtual scintillator rods in one layer.Alternatively, readout photodetectors (pixelated, strip, area) can becoupled to the 2D array of pixels or a combination of readoutphotodetectors can be coupled to one or more rods and pixels.

Yet another variation implements semi-continuous discrete or virtualscintillator 2D array of pixels in the top and bottom layers (pixels canbe aligned or offset between top and bottom layers). Improvements in oneor more detector resolution parameters (spatial, temporal, energy) canbe attained by sampling additional surfaces with readout photodetectors(including all or part of the face of a pixelated scintillator array)and/or by implementing encoding techniques such as implementingscintillators with spatially-varying parameters and/or applying patternsof WLS materials on at least one of the rod layer, the 2D pixel arraylayer.

Although examples of planar, square or rectangular sheets are shown inthe figures for discrete and semi-continuous structured 3D scintillatordetector modules other planar geometries (geometric and non-geometric)and non-planar geometries may be implemented. For example, a planar diskwith crossed rods can be implemented. Readout techniques can be modifiedto compensate for non-planar readout surfaces.

FIG. 13 illustrates a double-sided, semi-continuous rod-structured 3Dscintillator detector module 211 with a crossing angle of 90 degreescomprised of a single scintillator sheet (a double-sided semi-continuousstructured scintillator sheet with a rod structure) into which uniformand/or non-uniform physical gaps (cut/sawed/etched) are introduced withrespect to the top and bottom surfaces creating semi-continuous arraysof discrete rods 215, 216, respectively. Gaps can be unfilled,partially-filled or completely filled with one or more absorbing,scattering, reflective and WLS materials, including structuredmaterials. Alternatively, some or all of the physical gaps can bereplaced by virtual gaps (implemented using ion implantation,sub-surface laser engraving or other techniques known in the art).Variations in gap depth may be employed in order to modify (locally) thedistribution of fluorescence (or other optical) signals.

A variation of the double-sided, semi-continuous rod-structured 3Dscintillator detector module implements a single scintillator sheet (asingle-sided semi-continuous structured scintillator sheet with a rodstructure) into which physical gaps or virtual gaps (for example, viaion implantation) are introduced from only the top or bottom surface toa depth that is less than the thickness of the scintillator sheet andthereby defining two layers in the scintillator sheet (one structuredlayer and one continuous layer). The structured layer is comprised of anarray of semi-continuous discrete or virtual scintillator rods. Thisimplementation of a dual layer scintillator sheet is referred to as asingle-sided semi-continuous structured scintillator sheet with a rodstructure that can be employed as a single-sided semi-continuousrod-structured 3D scintillator detector module or with at least oneadditional scintillator layer to implement additional variation of 3Dscintillator detector modules.

As previously described herein, physical or virtual structures can beintroduced into (and on the surface of) one or more of thesemi-continuous scintillator rods (as well as the layer lacking an arrayof semi-continuous scintillator rods, for example, a 2D grid) in orderto improve at least one of spatial, temporal and energy resolution.Furthermore, previously herein-described encoding techniques can beemployed in order to enhance resolution. Readout photodetectors can bepositioned at one or both ends of the array of semi-continuousscintillator rods comprising either the top or bottom layer. Readoutresolution can be enhanced by implementing additional readoutphotodetectors positioned along one or more sides (including the face)of the scintillator sheet layer lacking an array of semi-continuousscintillator rods. The thicknesses of the top and bottom layers can bedifferent.

Although single-sided and double-sided semi-continuous structuredscintillator sheets are described as being comprised of two scintillatorlayers an equivalent description is a single scintillator layercomprised of two scintillator sub-layers. Double-sided semi-continuousstructured scintillator sheets with an intermediate layer can bedescribed as a single scintillator layer comprised of three scintillatorsub-layers. Thus, variations of a 3D scintillator detector module canimplement a single scintillator layer (a single scintillator sheet thatis comprised of at least two scintillator sub-layers).

Another variation on the single-sided, semi-continuous rod-structured 3Dscintillator detector module design is to implement a 2D array ofsemi-continuous discrete or virtual scintillator pixels on only the topor bottom layer (surface) of the scintillator sheet (a single-sidedsemi-continuous structured scintillator sheet with a pixel structure),forming a single-sided, semi-continuous pixel-structured 3D scintillatordetector module. Variations on this single-sided, semi-continuousrod-structured or pixel-structured 3D scintillator detector moduledesign is to implement a discrete scintillator 1D rod array or a 2Dpixel array layer coupled to a continuous or semi-continuousscintillator sheet layer. One readout configuration implements readoutphotodetectors (discrete, strip, area) on the surface of at least one ofthe pixelated layer, the continuous scintillator layer. Readoutresolution can be enhanced by implementing additional readoutphotodetectors positioned along one or more sides of the continuousscintillator sheet layer. As previously described herein, encodingtechniques can be employed in order to enhance resolution.

FIG. 14 illustrates a single-sided, semi-continuous rod-structured 3Dscintillator detector module 211 comprised of a single scintillatorsheet (a single-sided semi-continuous structured scintillator sheet witha rod structure) into which physical gaps 220 of uniform depth areintroduced in order to create an array of semi-continuous discretescintillator rods in only one (here, the top layer 301) of two layers.The bottom (second) layer 302 is shown as continuous with no gaps.(These layers can also be referred to as virtual layers since there isno discrete boundary to separate the two layers.) The depth of the top(first) layer, for example, is defined by the depths of the individualgaps. Gap depths can be uniform or non-uniform within the array ofsemi-continuous scintillator rods (variations in depth could beirregular or follow a geometric pattern such as a circular arc, aparabolic arc, an elliptic arc, a sinusoid, etc.). Variations in gapdepth may be employed in order to modify (locally) the distribution offluorescence (or other optical) signals.

Depending upon the application, the sheet can be inverted such that thetop and bottom layers reverse their positions with respect to thelocation of the radiation source. Alternatively, one or more physicalgaps can be replaced with virtual gaps creating one or more virtualscintillator rods. Readout photodetectors are positioned at both ends ofthe array of semi-continuous discrete or virtual scintillator rods. Forthe semi-continuous rod-structured or pixel-structured 3D scintillatordetector module described herein previously described encodingtechniques can be employed in order to enhance resolution.

The discrete structured 3D scintillator detector module and thedouble-sided, semi-continuous structured 3D scintillator detector moduleimplementations (including implementations previously described hereinsuch as arrays of crossed scintillator rods or an array of scintillatorrods with a pixelated scintillator array configurations) can be modifiedto incorporate at least one (discrete or continuous) intermediate layer.An intermediate layer can implement a different thickness than the topand bottom layers. The Nelson (U.S. Pat. No. 7,635,848) discretestructured 3D scintillator detector module implementation can bemodified to incorporate at least one discrete intermediate layer of rods(see Nelson, U.S. Pat. No. 8,017,906) or a transparent non-scintillatorlayer.

The scintillator or non-scintillator intermediate layer(s) can bestructured or unstructured (continuous). A transparent, structured orunstructured non-scintillator layer (as well as a scintillator layer)can be used to control the propagation of the fluorescence signalswithin and/or between top and bottom layers of scintillator rods. It canalso be used to modify the radiation field interacting with the bottomscintillator layer (as well as backscattered radiation from the bottomscintillator layer). The properties of a scintillator intermediate layer(material, fluorescence efficiency, temporal decay, etc.) can be thesame or be different from the properties of the top and bottomscintillator layers.

Thus, one variation implements an unstructured scintillator intermediatelayer (a continuous scintillator sheet). Another variation implements atleast one structured 2D pixelated scintillator intermediate layer (or ascintillator sheet with a 2D pixelated array). Yet another variationimplements at least one structured parallel rod array scintillatorintermediate layer (or a scintillator sheet with a parallel rod array).These parallel scintillator rods can be angled with respect to the topand bottom layers of crossed (arrays of) scintillator rods. Theintermediate layer(s) of parallel scintillator rods can implementphotodetectors at one end face or both end faces of the parallelscintillator rods or forgo photodetectors (reducing cost).

Yet a further variation implements a structured parallel rod arrayscintillator intermediate layer(s) (or a scintillator sheet with aparallel rod array) that is crossed with respect to the top and bottomlayers of arrays of scintillator rods (for example, the top and bottomlayers of arrays of scintillator are now parallel to each other if thereis one intermediate layer). The intermediate layer(s) of parallelscintillator rods (or a scintillator sheet with a parallel rod array)can implement photodetectors at one end or both end faces of theparallel array of scintillator rods.

Another variation with a discrete intermediate layer replaces one layerof discrete scintillator rods and the discrete intermediate layer withone single-sided semi-continuous structured scintillator sheet (whichimplements either a scintillator rod array structure and/or a pixelatedscintillator array structure). Yet another variation replaces theremaining layer of discrete scintillator rods with a second scintillatorsheet (with a rod and/or pixel structure) a single-sided semi-continuousstructured scintillator sheet with either a rod and/or pixel structure.

For the case of the double-sided, semi-continuous structured 3Dscintillator detector module with an intermediate layer the totalthicknesses of the semi-continuous discrete and/or virtual rods (orpixelated array) within the top and bottom layers of the double-sidedsemi-continuous structured scintillator sheet with a rod structureand/or a pixel structure) is now less than the thickness of thescintillator sheet. Previously herein-described physical or virtualstructures can be introduced into or onto one or more of the discreteand/or semi-continuous scintillator rods and/or a discrete intermediatelayer (for example, a 1D or 2D pattern of semi-continuous or virtualpixels can be introduced into discrete or virtual scintillator rods) ora continuous scintillator sheet intermediate layer in order to improvespatial and/or temporal and/or energy resolution. As previouslydescribed herein, encoding techniques can be employed in order toenhance resolution.

It should be understood, even if not explicitly stated, that thepreviously herein-described encoding techniques can be employed in orderto enhance resolution for the detector inventions described herein.Furthermore, continuous or discrete patterns of WLS materials can beimplemented with CT scintillator detectors (face-on or edge-on) as wellas scintillator and non-scintillator materials in which Cherenkovradiation is generated and thereby provide at least one of: spatialencoding, wave length shifting, spatial redistribution ofshifted-fluorescence signals, temporal shifting.

The WLS materials can be applied to one or more surfaces of scintillatorrods or pixels used for CT x-ray detectors. For example, a scintillatorwith an emission spectrum that offers a poor optical match to thespectral response of a photodetector may be suitable for x-ray and/orgamma ray imaging with the application on appropriate scintillatorsurfaces of one or more WLS materials of a more desirable emissionspectrum(s). In a similar manner the detection of short-wavelengthoptical Cherenkov radiation signals can also benefit from theapplication of WLS materials.

Temporal shifting (and wavelength shifting) can be used as a techniquefor spatial encoding for scintillator fluorescence radiation as well asCherenkov radiation. The spatial redistribution of shifted-fluorescencesignals can be used to reduce light trapping of a fraction of thescintillator fluorescence signal (or Cherenkov radiation) emitted withinthe scintillator volume, including optical fibers, scintillator fibers,face-on and edge-on scintillator detectors employed in CT, SPECT and PETdetectors (or combinations thereof).

Readout configurations can vary according to resolution (spatial,energy, temporal) requirements. In general, photodetector readoutdensities for pixelated arrays (as well as block, slab, etc.scintillator geometries) can vary from one photodetector perscintillator pixel to less expensive implementations employing sparsearrays (often with light guides). In many instances a crossed rodgeometry typically employs readout photodetectors at one end of each ofthe scintillator rod arrays whereas a geometry that employs only asingle array of scintillator rods typically employs readoutphotodetectors at both ends of the scintillator rod array. (Note thatpreviously herein-described variations include a readout at one end withor without encoding techniques.) Some applications may requireimprovements (or can tolerate reductions) in one or more resolutionparameters. If resolution improvements are required for either the oneend or both ends readout configuration then additional surfaces can besampled with readout photodetectors (including all or parts of the endsof one or more scintillator rod arrays, all or parts of one or moresides of discrete or continuous intermediate layers, all or part of theface of a pixelated scintillator array) and/or encoding techniques canbe implemented.

FIG. 15A illustrates a discrete structured 3D scintillator detectormodule 221 with crossed scintillator rods 225 implementing a discreteintermediate layer 315.

FIG. 15B illustrates a double-sided, semi-continuous structured 3Dscintillator detector module 221 with crossed scintillator rods 227implementing a continuous intermediate layer 320 (a double-sidedsemi-continuous structured scintillator sheet with a rod structure andan intermediate layer). An alternative implementation employs a virtualpixelated intermediate layer.

FIG. 15C illustrates a double-sided, semi-continuous structured 3Dscintillator detector module 221 with parallel scintillator rodsimplementing a continuous intermediate layer. Alternativeimplementations employ a virtual pixelated or crossed scintillator rodsintermediate layer.

A conventional 2D scintillator rod parallel array detector module(face-on geometry) employs an array of discrete scintillator rods with aphotodetector readout on one end (the bottom side or bottom end) of thearray. Spatial resolution can be increased from 2D to 3D (addingdepth-of-interaction (DOI) resolution for a face-on geometry orsub-aperture resolution (SAR) for the edge-on geometry) by positioningphotodetectors (at least one photodetector may implement 2D spatialresolution) at both ends of the scintillator parallel array (usingsignal weighting between opposite photodetectors at the ends of ascintillator rod to compare signal strength at both ends) with reducedrequirements for coincidence measurements between photodetectors at thetwo ends of the rods.

As previously described herein, weighting can also be employed with TOFresolution for a photodetector at both ends of a scintillator rod array(comprised of discrete rods or rods within a scintillator sheet or acombination of discrete rods and rods within a scintillator sheet).Alternatively, photodetectors with TOF resolution can be employed atonly one end of a scintillator rod array using a comparison between theinitial pulse and a direct reflected pulse or an effective reflectedpulse (a delayed signal) from the opposite end (for example, a WLSreflected pulse). One end (one-sided) and both end (two-sided) readoutconfigurations can take advantage of previously described encodingtechniques in order to implement 3D detector modules. For example, oneor more scintillator properties of a scintillator rod such as temporaldecay distribution, spectral distribution, conversion efficiency can bevaried spatially along the length of a scintillator rod according to anencoded pattern. (Spatially-varying scintillator properties can also beused as a technique for encoding scintillator sheets.) Another techniquefor improving TOF resolution is to utilize the gamma ray interactionpositional information along the length of the scintillator rod tocorrect for the position-dependent propagation time of the opticalsignal.

Continuous or discrete (encoded) patterns of WLS materials can beapplied along the length of a rod such that wavelength and/or pulseshape properties vary with position. Both encoding techniques can beused together. (If the area of a rod face is sufficiently large 3Dresolution within the rod is possible.) The top side of the 2D face-ondetector design (one end readout) could employ a reflector (focused orplanar) or one or more WLS materials (with or without a reflector). A 3Dface-on detector design (or a 3D edge-on detector design) employing aparallel array of scintillator rods with a photodetector readout at thetop and bottom (or both sides for the edge-on geometry) can implementthese encoding techniques described for the 2D face-on or edge-ondetectors. Potential benefits include enhanced DOI spatial resolution(sub-aperture resolution (SAR) for the edge-on geometry), energyresolution and reduced ambiguity.

Two or more scintillator rods (including all rods) within a single rowor within multiple rows of the parallel array of scintillator rods canbe replaced by scintillator sheets (including structured scintillatorsheets with a rod structure, continuous scintillator sheets, singlesided and double-sided semi-continuous structured scintillator sheetswith a rod structure). The encoding techniques described herein for usewith scintillator rods can be implemented with scintillator sheets.

It should be noted that 2D/3D edge-on and face-on readout geometries arenot limited to the ends of scintillator rods. 2D/3D edge-on and face-on(or combinations of edge-on and face-on) readout geometries alsoinclude, but are not limited to, the sides of scintillator rods, thesides and ends of scintillator rods, one or more sides of a scintillatorsheet (including thick scintillator sheets/blocks).

The discrete structured 3D scintillator detector module design, as wellas the variations taught herein, have described planar geometryimplementations. There are radiation imaging applications (medicalimaging, inspection, security, astronomy, high energy physics, etc.)that would benefit from the implementation of a focused detectorgeometry such as focused discrete structured or semi-continuousstructured 3D scintillator detector modules (which can be employed inface-on or edge-on orientation). Consider the example of a face-onorientation of focused 3D scintillator detectors suitable for deploymentin a spherical/cylindrical imaging geometry (other geometries can beimplemented).

The focused discrete structured 3D scintillator detectors can beimplemented by using crossed top and bottom spherical/cylindrical shelllayers of parallel arrays of discrete and/or virtual scintillator rodsin which the parallel rods in each layer arecut/sawed/etched/ion-implanted/laser-engraved using a scintillatorspherical/cylindrical shell (sheet) segment. In a similar manner,focused double-sided semi-continuous structured 3D scintillatordetectors can be implemented by using a single spherical/cylindricalshell layer and forming parallel arrays of semi-continuous discreteand/or virtual scintillator rods. All of the 3D detector moduleimplementations previously discussed herein (double sided orsingle-sided, semi-continuous structured 3D scintillator detectormodules, discrete structured 3D scintillator detector modules andsemi-continuous structured 3D scintillator detector modules with atleast one intermediate layer, etc.) can be implemented with a focusedgeometry.

An alternative implementation is to approximate the shape ofscintillator spherical/cylindrical shell segments with planar segmentsin which the arrays of discrete and/or virtual scintillator rods (orsemi-continuous discrete and/or virtual scintillator rods) follow thecurvature of the spherical/cylindrical geometry. The edges of the planarsegments can be tapered if desired to form a close fit with nearestneighbor planar segments.

Note that in the case of a face-on or edge-on structured 3D scintillatordetectors employed in a cylindrical geometry focusing in the radialdirection is beneficial whereas focusing in the axial direction may beof reduced benefit in some applications relative to the cost ofimplementation. The benefit of axial focusing (see FIG. 3, FIG. 11) isreadily apparent when point-like or relatively small radiation sourcesare used for imaging (e.g. x-ray radiography applications including, butnot limited to, tomosynthesis, CT, cone beam CT, chest x-ray). Focusingcan be implemented in PET, SPECT, probes/mini-PET or mini-Gamma cameras.An alternative for PET imaging is to implement only a subset of 3Ddetectors to implement axial focusing, thus reduced cost while allowingimproved imaging over smaller regions of interest.

Focusing can be implemented with face-on structured mold detectors(semiconductor, scintillator, gas, etc.) used in single layer ormultiple layer flat panel ionizing radiation detector applications(e.g., chest x-ray, cone beam CT, tomosynthesis, etc.) by varying theslant angle that holes (containing the detector material) make withrespect to the flat panel surface. When used with a point-like radiationsource the slant angle of the holes typically increases away from theflat panel center in order to compensate for the divergence of thepoint-like radiation source at the location of the flat panel.Optionally, for implementations using semiconductor materials the anodes(and/or cathodes) can be segmented (providing energy resolution and/orDOI capability for a face-on orientation and SAR for an edge-onorientation). Anode and cathode slant angles can be implemented withstructured 3D silicon, 3D GaAs, 3D CdTe, 3D CZT, 3D Ge, 3D GaP, 3D GaSe,3D diamond, 3D Se, 3D PbS, 3D InP, 3D PbI₂, 3D HgI₂, 3D PbO, 3D CdS, 3DTlBr, etc. (including doped semiconductors and other dopedimplementations of these materials).

FIG. 16 illustrates a focused discrete structured 3D scintillatordetector module 230 with crossed top and bottom cylindrical shell layersof parallel arrays of discrete scintillator rods 231, 232. Othergeometries (including, but not limited to, spherical layers) can also beimplemented depending on the application. One or more discretescintillator rods can be replaced with virtual scintillator rods. The 3Ddetector module implementations previously discussed herein can beimplemented with a focused geometry, including implementations withintermediate layers.

3D Scintillator and Optical Fiber Detectors

Another implementation is a discrete structured 3D scintillator fiberdetector module that implements one or more independent layers of 1Dparallel arrays of scintillator fibers and/or at least one set ofcrossed coupled top and bottom (planar, spherical, cylindrical, etc.)layers of 1D parallel arrays of scintillator fibers (previouslyherein-described encoding techniques can be employed in order to enhanceresolution. including spatially-varying the properties of thescintillator fiber materials and the use of WLS materials and patternedWLS materials). (Note that these encoding techniques can also beemployed with 2D arrays of scintillator fibers.)

One of the crossed scintillator fiber layers can be replaced by anoptical fiber layer (which can implement WLS materials and patterned WLSmaterials). Spatial positioning within scintillator fiber (or fiberoptic) layers can be determined by implementing one or more encodingtechniques and/or employing photodetectors at both ends (using signalweighting to compare signal strength at both ends and/or TOF resolutiontechniques to compare initial pulse arrival times at both ends) or atone end (using TOF resolution to compare arrival times of the initialpulse with the direct reflected pulse or an effective reflected pulse (adelayed signal) from the opposite end (such as a WLS reflected pulse) ofthe independent layers (this is also applicable for 2D arrays ofscintillator fibers, layers or arrays of non-scintillator fibers,crossed layers of scintillator fibers, crossed layer of non-scintillatorfibers).

One or more non-scintillator supporting layers (as previously describedherein) can be employed. The properties of the scintillator fibers(temporal decay, fluorescence efficiency, spectral distribution,energy-dependent and/or particle-dependent stopping power) within aparallel array as well as between crossed parallel arrays can beselected based on the application. For example, a fast scintillatorfiber may be used for TOF applications. Previously herein-describedreadout geometries (photodetectors at one end or both ends) can beimplemented for each layer of scintillator fibers (or optical fibers).

A variation of the discrete structured 3D scintillator fiber detectormodule enhances detection efficiency by incorporating additional sheet,2D or 3D scintillator (or semiconductor) detectors (which may alsofunction as converters) which can be positioned behind, in front of, orbehind and in front of the discrete structured 3D scintillator fiberdetector module (depending on the application) in a stackedconfiguration. A discrete structured 3D scintillator fiber detectormodule can be manufactured with an area comparable to or greater than atleast one sheet, 2D, or 3D scintillator detector (it can span multipledetectors).

A scintillator layer or layers (structured or continuous, comprised ofdiscrete scintillator rods/pixels and/or scintillator sheets) can becoupled to one or more scintillator or optical fiber layers. Ascintillator layer can be positioned between and coupled to crossedlayers or parallel layers of scintillator fibers or optical fibers. Ascintillator fiber or optical fiber layer can be shared between (top andbottom) scintillator layers and a scintillator or optical fiber layercan be coupled to and span multiple scintillator layers. The structuredor continuous scintillator layer(s) can employ one or morephotodetectors on one or more sides to enhance resolution.

If the scintillator or optical fibers are transparent to thescintillator detector fluorescence (or employ WLS materials to convertscintillator fluorescence to wavelengths that can be transmitted throughthe scintillator or optical fibers) then photodetectors can be coupledto scintillator layer sides coupled to the scintillator or opticalfibers. Detection efficiency gains can be realized by incorporatingadditional sheet, 2D or 3D scintillator (or semiconductor) detectors(which may also function as converters). For example, a simple areaphotodetector (providing energy resolution and/or temporal resolution)could be coupled to one surface (crossing the Z axis) of a scintillatorsheet/block (or an array of uniform or non-uniform scintillator sheetswhich may be optically-coupled, partially coupled, oroptically-isolated) to provide energy resolution while crossed layers ofscintillator or optical fibers (or a combination of both) are coupled totwo surfaces (crossing an X axis and a Y axis).

One or both layers of scintillator fiber or optical fibers (or acombination of both) can be coupled to one or more scintillatorsheets/blocks. Improvements in spatial resolution and/or a reduction inthe number of scintillator and/or optical fiber layers can be achievedby replacing the simple area photodetector with an area detector with areadout at its four corners, a strip array photodetector, a 2D arrayphotodetector.

A scintillator fiber layer (or an optical fiber layer) can be sharedbetween (or coupled to) two adjacent intermediate scintillator layerswithin a stack geometry or other configurations. For example, theintermediate layers could be arrays of scintillator rods in parallel ina stack configurations. In another example, intermediate layers ofsheets/blocks could share a scintillator fiber layer (or an opticalfiber layer) in a stack geometry or adjacent sheets/blocks could share ascintillator fiber layer (optical fiber layer) in a differentconfiguration (or one scintillator fiber layer (optical fiber layer) canbe shared between two layers of sheets/blocks while the otherscintillator fiber layer (optical fiber layer) is shared between twoadjacent sheets/blocks. A scintillator fiber layer (optical fiber layer)can span one or more sheets/blocks. The structured or unstructuredscintillator detector(s) can employ one or more photodetectors on one ormore sides not coupled to the scintillator fibers in order to enhanceresolution.

The choice of a scintillator fiber core is not limited to glass andplastic scintillators (including embedded nanoparticles), scintillatorfiber cores can also be made from scintillator materials. Furthermore,by varying one or more scintillator properties of the scintillator fibercore (temporal decay distribution, spectral distribution, conversionefficiency, energy-dependent and/or particle-dependent stopping power)along its length according to a predetermined pattern then positionalinformation can be encoded into the scintillator fiber core. Patternscan also be imposed using WLS materials.

Implementations of encoded scintillator fiber cores may be used toenhance resolution for the case of crossed fiber layers (with or withoutan intermediate scintillator detector layer) or (if encoding permitssufficient accuracy) alternatively allow the use of uncrossed fiberlayers. Cherenkov radiation can also be recorded. Readout photodetectorsare coupled to one end or both ends of each array of crossedscintillator fibers. If the readout photodetectors are only coupled toone end then the opposite end is typically covered with a reflectormaterial and/or a WLS material. This represents an alternative toreading out both ends of each single layer of scintillator fibers.

A variation of the discrete structured 3D scintillator fiber detectormodule coupled to at least one intermediate scintillator layer(structured or unstructured, including scintillator rods and/or pixels,scintillator sheets/blocks, continuous scintillator sheets) implements alayer of optical fibers (including WLS materials) in place of one orboth layers of scintillator fibers that are cross couple to the at leastone intermediate scintillator layer (a discrete structured 3Dscintillator fiber-fiber optic detector or a discrete structured 3Dfiber optic-fiber optic detector, respectively). Each layer of opticalfibers is separated from its nearest crossed neighbor layer ofscintillator fibers or optical fibers by an intermediate scintillatordetector layer.

These scintillator and/or fiber optic layers can span multipleintermediate scintillator detectors or a single intermediatescintillator detector. A scintillator fiber layer or a fiber optic layercan be shared between (coupled to) two adjacent intermediate (discreteor continuous) scintillator layers within a stack configuration. Thestructured or unstructured scintillator detector(s) can employ one ormore photodetectors on one or more sides not coupled to the scintillatorfiber layer or the optical fiber layer in order to enhance resolution(at additional cost).

FIG. 17 illustrates a discrete structured 3D scintillator fiber detectormodule 240 in which a structured intermediate scintillator layer(comprised of multiple scintillator detector sheets/blocks 345) ispositioned between and coupled to crossed planar layers of scintillatorfibers 241, 242. A scintillator fiber layer (or, alternatively a fiberoptic layer) is not constrained to terminate at the outer boundary of astructured intermediate scintillator layer (as shown in FIG. 17) or anunstructured intermediate scintillator layer.

It should be noted that the 3D particle detectors that implementlight-sharing designs may have both positive benefits and negativepotential consequences. Positive benefits, depending on the design,typically include features such as cost-savings and detector designflexibility. Negative potential consequences, depending on the design,typically impact one or more features such as: detection rates, energyresolution, temporal resolution, coincidence detection.

Consider the following example of a double-sided semi-continuousstructured 3D scintillator detector module. A 4 mm thick, square-shaped(other geometric shapes such as rectangles, circles, etc. can beimplemented according to detector requirements) scintillator sheet (adouble-sided semi-continuous structured scintillator sheet with a rodstructure) has a parallel array of semi-continuous discrete scintillatorrods, each 2 mm deep and 2 mm wide, cut/sawed/etched/ion-implanted intothe top surface. An identical parallel array of semi-continuous discretescintillator rods (2 mm deep, 2 mm wide) iscut/sawed/etched/ion-implanted/laser-engraved into the bottom surfacebut in a crossed direction (typically 90 degrees) to the orientation ofthe top surface array of rods.

The ends of the cut/sawed/etched/ion-implanted/laser-engravedsemi-continuous discrete scintillator rods can be shaped, patterned,coated, etc. as desired to provide good optical coupling tophotodetectors. One or more coatings can be applied to any accessiblesemi-continuous discrete scintillator rod walls (specular reflective,diffuse reflective, low index, WLS, etc.). Coatings can be continuous orapplied in patterns. Accessible semi-continuous discrete scintillatorrod walls can be completely or partially polished (smooth), etched orroughened (patterns can be implemented). Internal structures can beintroduced into the scintillator sheet. Previously herein-describedencoding techniques (including varying one or more scintillatorproperties of a scintillator rod along its length) can be employed inorder to enhance resolution. Readout photodetectors are coupled to atleast one end of at least one array of scintillator rods comprising thetop and bottom layers.

Discrete and Semi-Continuous Structured 3D Scintillator Detectors withIntermediate Layers

Nelson (U.S. Pat. No. 7,635,848) has previously taught theimplementation of at least one additional intermediate layer of crossedarrays of discrete scintillator rods. For the case of one intermediatelayer the discrete scintillator rod arrays of the top and bottom layersare parallel to each other and crossed with respect to the discrete rodarray of the one intermediate layer. An alternative is to replace the atleast one additional intermediate layer of an array of discretescintillator rods with at least one non-scintillator or scintillatorsheet (or spacer).

Consider the case of a scintillator sheet intermediate layer in whichthe intermediate layer scintillator material can be different from thescintillator materials implemented in the top and bottom layers.Coatings (including structured/patterned coatings, WLS coatings,coupling materials, etc.) can be added to the surfaces and internalstructures can be introduced into the intermediate scintillator sheetlayer (as well as the top and bottom layers) as needed to enhanceoptical coupling and/or to direct the propagation of optical signals(fluorescence and in some instances Cherenkov radiation). Previouslyherein-described encoding techniques can be employed in order to enhanceresolution for the intermediate scintillator sheet layer as well as thetop and bottom layers.

The thickness of the intermediate layer of scintillator material istypically chosen to be one times the thickness of the top or bottomlayer although other multiples greater than or less than one may also bechosen as advantageous based on a factors such as energy resolution,temporal resolution, spatial resolution, stopping power, cost ofmaterials and cost of readout electronics. In one implementation thescintillator sheet intermediate layer implements a discrete or virtualarray of rods or pixels. In another implementation the scintillatorsheet intermediate layer is continuous. In yet another implementationthe scintillator sheet intermediate layer is semi-continuous(single-sided, double-sided, double-sided with an intermediate layer).

The result of adding an intermediate layer is that in someimplementations the readout optical signals will be more spread-out(dispersed). If spreading out of optical signals is determined to bedesirable independent of using an intermediate scintillator layer thenan alternative is to implement a transparent non-scintillator sheet(fiber array, plastic, epoxy, gel, fluid, air) which may have aninternal structure. The thickness of the transparent non-scintillatorintermediate layer can be selected to control the spread of the opticalsignal. The index of refraction (IOR) of the non-scintillatorintermediate layer material can be selected (within limits) to improveat least one of spatial resolution, temporal resolution, optical signalcollection. The non-scintillator intermediate layer can have an internalstructure including patterns of embedded microstructures such as low IORmicrospheres (or microellipsoids, etc.) for the purpose of redirectingoptical signals propagating between the top and bottom layers.

Consider the example of a semi-continuous, double-sided with anintermediate layer structured 3D scintillator detector module whereinthe scintillator sheet/block has dimensions of W=L=N1 mm, thickness=6mm. (In this example assume N1 is an integer multiple of 2 forconvenience.) An array of (N1×0.5) semi-continuous discrete or virtualscintillator rods with dimensions of 2 mm×N1 mm and a consistent depth(thickness) of 2 mm are cut/sawed/etched/ion-implanted/laser-engravedinto each of the top and bottom surface. (Note that othersemi-continuous discrete or virtual rod depths can be implemented tomodify light propagation and the rod depth can be varied spatiallybetween rods and along the length of a rod.) For the case of aconsistent rod depth of 2 mm the effective intermediate layer has aconsistent effective thickness of 2 mm (other effective thickness can beimplemented). In some implementations the effective intermediate layercan be readout from one or more sides or not at all (relying exclusivelyon the readout detectors coupled to at least one end of at least onearray of semi-continuous discrete scintillator rods comprising the topand bottom layers. The result of adding an intermediate layer is thatthe readout optical signals will generally be more dispersed.

Detector Stack Configurations

The Nelson (U.S. Pat. No. 7,635,848) discrete structured 3D scintillatordetector module designs as well as the variations taught herein can beimplemented in a detector configuration with a depth of 1 module or adepth of at least 2 modules (a stack). A stack configuration of 3Ddetector modules extends the flexibility inherent to a single moduleimplementation. Properties of modules within a stack can range frombeing identical to being highly dissimilar. Individual layer thickness,module thickness, individual layer spatial resolution, individual layerscintillator material(s) as well as optional conversion materials can betailored for specific imaging (and/or tracking) applications includingsingle and multiple radiation fields encountered in x-ray radiographyand CT imaging, nuclear medicine and PET imaging, radiation therapy (ofall types), probes, high energy physics, astronomy, industrial imaging,homeland security, etc.

The stack configuration offers significant flexibility in the choice ofdetectors and therefor is not limited to implementing only structured 3Dscintillator detector modules. Face-on and edge-on scintillator and/orsemiconductor detectors (included structured scintillator andsemiconductor detectors as well as scintillating fibers or WLS fibers)can be incorporated into a stack configuration. A stack configurationmay enable the use of scintillator materials with one or more propertiessuch as stopping power, fluorescence efficiency, temporal decay,manufacturability that would otherwise limit their use in a conventionalPET detector implementation. Additional variations on the 3Dscintillator detector module stack introduce at least one semiconductoror structured detector layer between at least two adjacent 3Dscintillator detector modules within a stack and/or replace at least one3D scintillator detector module with at least one semiconductor orstructured detector layer within a stack.

For example, consider the case of a detector stack implemented for SPECTand PET imaging. The at least one front-end structured 3D scintillatordetector module in the stack can be adapted for low energies such as140.5 keV (Tc-99m) used in SPECT imaging while the back-end structured3D scintillator detector module(s) can be adapted for PET imaging.Slower scintillators may be acceptable for SPECT whereas fasterscintillators would be preferred for PET. Alternatively, the same or adifferent PET scintillator may be employed for SPECT imaging.

Furthermore, consider the case of a detector stack implemented for CT(ring CT, cone beam CT, tomosynthesis) and PET imaging. The front-enddetector module(s) could implement an edge-on or face-on scintillator orsemiconductor detector module(s) offering at least one of energyintegration, photon counting, photon counting with energy resolutioncapability suitable for CT (including tomosynthesis). Optionalenergy-dependent radiation attenuation filter layers can be insertedbetween one or more layers within the front-end modules(s) and/or anenergy-dependent radiation attenuation filter layer can be insertedbetween the front-end module(s) and the back-end module(s) to modify thespatially-dependent radiation spectrum (non-scattered and scatteredphotons, characteristic x-rays, bremsstrahlung, electrons) in order toenhance energy resolution and/or reduce the detection of undesirableradiation. If photon counting with energy resolution is implemented andexcessive count rates that significantly degrade energy resolution areexperienced then the electronics can automatically implement eitherphoton counting or energy integration capability. The back-end detectormodule(s) could implement optional edge-on or face-on scintillator orsemiconductor detector modules including structured 3D scintillatordetector modules.

Multi-Energy CT, CT with PET/SPECT

Consider the case of a planar or focused edge-on detector array moduleimplemented for CT (ring CT, cone beam CT, tomosynthesis) with orwithout a stack detector implemented for PET (or SPECT) imaging. Anedge-on, n-level multispectral CT detector module (aligned with theaxial direction) implements one or more CT scintillators for the nlevels with photodetector readouts. Optional energy-dependent radiationattenuation filter layers with can be inserted between one or more CTscintillator layers. If an additional back-end detector module (forexample, a SPECT and/or PET detector module), separate or integrateddirectly with the CT detector module, is employed behind the n-levelmultispectral CT detector module then an optional energy-dependentradiation attenuation layer can be inserted between the CT detectormodule and the additional back-end detector module.

The choice of detector element scintillator and/or the depth of thescintillator (in the radial direction) can be adjusted to compensate forx-ray beam hardening as it propagates through each of the n levels ofscintillator material (and any optional attenuation filters),constituting a “poor man's” n-level multispectral CT scintillatordetector (also applicable for cone beam CT and tomosynthesis). Energycalibration will be implemented.

For the case of n=1 there is only one energy level (conventional, singlespectrum CT detector). A simple example of a multispectral CT detectoris the n=2 case (dual energy). The top scintillator level (level 1)would implement a low-Z x-ray/gamma ray scintillator (or a reducedthickness of a high-Z x-ray/gamma ray scintillator) and the bottomscintillator level (level 2) would implement a high-Z x-ray/gamma rayscintillator with 2 levels of photodetectors providing signal readout.An x-ray energy filter can be positioned between the two levels ofscintillator and optionally in front of the top level (level 1)scintillator.

Each two-level edge-on detector array is aligned with the axialdirection and distributed in a circle or semi-circle (full ring orpartial ring CT detectors, respectively). Similar dual energy(multi-energy) configurations can be implemented for planar cone beam CTand tomosynthesis detectors (optionally, focusing can be implemented inthe radial and/or axial direction) as well as curved cone beam andtomosynthesis detector implementations. The distribution of scintillatorlevels and/or scintillator materials can be varied in the axialdirection to accommodate different x-rays spectrums (from a singlefiltered x-ray source or from multiple x-ray sources distributed in theaxial direction). This variation along the axial direction can also beimplemented if an n-level multispectral CT semiconductor detector isemployed. The spectral CT designs described herein may also beimplemented for cone beam CT and tomosynthesis.

The flexibility of the multilayer design includes the use of edge-onand/or face on detectors, one or more detector materials, at least oneof energy resolution, PC capability, energy integration capability. Somedesigns involve sharing one or more layers of detectors for two or moreimaging modalities (CT, SPECT, PET, Compton). Although conventional CTbenefits from the use of a point-like x-ray source (with implieddirectionality) the typical broad x-ray energy spectrum and high x-raycount rates per pixel may impose a financial and engineering burden onthe multi-spectral CT designers.

A variation of the previously herein-described “poor man's” n-levelmultispectral CT scintillator detector sacrifices the x-ray stoppingpower of the layer 1 scintillator with a higher quality (such as asemiconductor) energy-resolving detector layer(s) while emphasizing thestopping power of the energy-integrating detector layer(s) (which may beshared, for example, with a PET detector system). A thinner version of adetector (with respect to the direction of the incident radiation)typically offer lower stopping power but with potential advantages suchas faster response, better energy resolution, a lower count rate and alower cost.

For example, in one implementation, a dual layer spectral CT detectoremploys a first layer comprised of relatively thin (in terms of x-raystopping power) pixelated, face-on semiconductor detectors and thesecond layer implements a relatively thick (in terms of x-ray stoppingpower) pixelated, face-on scintillator detector. The first layer offerssufficient stopping power to ascertain the incident x-ray spectraldistribution (within the guidelines established for patient radiation)while limiting pulse pile up effects. The second layer (which could, forexample be based on a pixelated, face-on CT detector array) offerssubstantial stopping power and functions as an energy integrator.

The total detector energy for pairs (or multiple pairs) of aligneddetector pixels can be summed. Spectral data from individual orneighboring detector pixels of the first layer can be evaluated.Corrections can be applied to account for scattering within and betweendetector layers as well as the responses of the first and second layerdetectors. The spectral and total detector energy data can then beevaluated with respect to calibration data in order to implementmulti-spectral CT reconstruction. Pixels sizes of the first and secondlayer detectors can be the same or they can be different (as previouslydescribed herein for the case of CT-PET).

Other variations of this two layer spectral CT design employ edge-ondetectors in the first and second layers or a combination of edge-on andface-on detectors in the two layers. The first layer is not limited tosemiconductor detectors. One alternative is structured semiconductordetectors. The second layer is not limited to scintillators.Semiconductor or structured semiconductor detectors (or other detectortypes) can be employed.

This multi-spectral CT design can be extended to three or more layersincluding one or more relatively thin layers (multiple thin layers canuse the same or different semiconductor or structured semiconductordetector materials) followed by one or more relatively thick layers(which can be of the same or different scintillator, semiconductor,structured semiconductor, etc.) detector materials. For example, one ormore relatively thin semiconductor detector layers could be combined (asthe top layer or layers) with a “poor man's” n-level multispectral CTscintillator detector. This design may also be employed with cone beamCT detectors and tomosynthesis detectors.

Multispectral CT detectors (including ring CT, cone beam CT andtomosynthesis detectors) can be implemented in scanner systems with oneor more x-ray sources (x-ray tubes, scanning electron beams) and thex-ray sources can be operated at one or more (switched) voltage levelsproviding one or more x-ray spectrums. Ring CT geometries include fullring and partial ring configurations. A partial ring CT geometry employsat least one partial ring and at least one x-ray source. A full ring CTgeometry typically employs multiple x-ray sources or a scanning electronbeam source, permitting faster acquisition times with reduced mechanicalwear.

CT data (as well as PET data) can be correlated with other patientmonitoring data (such as EKG data) for purposes of gating. CT scanners(and tomosynthesis scanners) typically include at least one ofcollimators and/or numerical techniques to provide x-ray scatterreduction and/or to correct for scatter. An external collimator can bemade mobile if desired. The edge-on detector design will typicallyinclude external collimation and may incorporate internal collimation toreduce x-ray cross talk and/or shield readout ASICs from directradiation.

FIG. 18A illustrates an edge-on, planar n-level (n=3) multispectral CT(ring CT, cone beam CT, tomosynthesis) scintillator detector 501comprised of an array of detector elements which implementphotodetectors 281-283 coupled to the side faces of scintillators271-273 with incident radiation 109. FIG. 18A illustrates only 3detector elements in each row for illustrative purposes. The number ofdetector elements per row can greatly exceed 3 for many applications.Detector element scintillator materials and/or depths can be the same orthey can be different depending on the application.

FIG. 18B illustrates an edge-on, planar n-level (n=4) multispectral CT(ring CT, cone beam CT, tomosynthesis) scintillator detector 501comprised of an array of detector elements which implementphotodetectors 281-284 coupled to the side faces of scintillators271-274 with incident radiation 109. For each value of n (1-4) thescintillator materials and/or scintillator depths can be the same ordifferent. The value of n within each column can be same or can bevaried along the direction of the rows of detector elements. An x-rayenergy filter can optionally be positioned between one or more adjacentlevels of scintillators. Furthermore an energy filter can optionally bepositioned in front of the top level (level 1) scintillator.

Although FIG. 18A illustrates individual rows of uniform detectorelements (along the axial direction) the flexibility of this designpermits n to be varied along the axial direction according to theapplication requirements. One implementation butts independent edge-on,planar multispectral CT detectors with different values of n in theaxial direction. Another implementation involves electronically (orduring post-processing) combining two or more adjacent photodetectors(pixel merging) and thereby reduces the local value of n. For example,the n=4 multispectral CT (ring or cone beam CT, tomosynthesis) detectorelements could be positioned to cover the volume of the patientrequiring maximum spectral resolution whereas other patient volumescould be covered by n<4 multispectral CT (or tomosynthesis) detectorelements.

The edge-on, n-level scintillator detector (or a section thereof) can bereplaced by an edge-on, n-level semiconductor detector (including astructured semiconductor detector). Furthermore, one or more edge-onscintillator detector levels can be replace by edge-on or face-onsemiconductor (or structured semiconductor) detector levels. Optionally,at least one energy-dependent radiation attenuation filter can beincluded with the edge-on, n-level semiconductor detector (with orwithin the structured semiconductor detector) or combinations of edge-onscintillator detector layers and semiconductor (structuredsemiconductor) detector layers.

Continuous or discrete patterns of WLS materials can be implemented withface-on scintillator CT detectors and edge-on (single energy ormultispectral) scintillator CT detectors described herein, providing atleast one of spatial encoding, spatial redistribution ofshifted-fluorescence signals, wavelength shifting, temporal shifting indedicated CT scanners or scanners in which CT is integrated with atleast one of a SPECT, PET or Compton camera. For example (as previouslydescribed herein), a scintillator with an emission spectrum that offersa poor optical match to the spectral response of a photodetector may besuitable for x-ray and/or gamma ray imaging with the application onappropriate surfaces of one or more WLS materials of a more desirableemission spectrum(s). The spatial redistribution of shifted-fluorescencesignals can be used to reduce light trapping of a fraction of thescintillator fluorescence signal emitted within the scintillator volume.

The n-level, multispectral CT (cone beam CT, tomosynthesis) detector canbe integrated with a PET and/or SPECT and/or Compton detector (theback-end detector) implementing at least one of simultaneous orsequential acquisition of CT and PET and/or SPECT images. One ormultiple x-ray sources (x-ray tubes, scanning electron beams, etc.) canbe employed. Multiple x-ray sources can be distributed along thecircumference of a ring and/or along the axial direction. Multiple x-raysources can be used to provide at least one of: reduced acquisitiontimes (improved temporal resolution), multiple x-ray spectrums.

In general, if gaps within the PET and/or SPECT and/or Compton detectorsare due to the presence of one or more x-ray sources then more completesampling can be achieved by inserting removable PET and/or SPECTdetector modules (electronically linked to the existing detectormodules) into the gaps when the x-ray sources are no longer needed orexisting PET detector modules can be temporarily rotated into theposition of the gaps. The extent of the CT detector and PET and/or SPECTand/or Compton detector along the circumference or in the axialdirection can be the same or different. If a partial ring geometry CTimplementation is preferred to a full ring geometry implementation thenthe properties of any external PET and/or SPECT and/or Compton detectorsused to implement a functional PET and/or SPECT scanner can be the sameor different from the properties of the integrated CT-PET and/or SPECTdetectors.

For example, the external PET and/or SPECT and/or Compton detectorscould employ non-functional substitutes for front end CT detectormaterials employed with the integrated CT-PET and/or SPECT detectors inorder to reproduce the scatter and absorption properties of the frontend CT detector. Similar principles apply for the cases of cone beam CTand tomosynthesis imaging detectors. Previously herein-described energycalibration techniques can be employed. Positioning stacked structured3D scintillator detectors (or other 2D/3D PET and/or SPECT and/orCompton detectors including edge on and/or face on high-Z scintillatoror semiconductor detectors or structured semiconductor detectors)beneath this edge-on, n-level multispectral CT detector for PET and/orSPECT imaging is one option.

Another option is to extend the edge-on, n-level multispectral CTscintillator detector in depth to include m-levels of edge-on PET and/orSPECT scintillator(s) detectors (edge-on, n/m-level CT-PET, CT-SPECT orCT-SPECT-PET detector) in place of the 2D or 3D scintillator detectors,etc. The dimensions of the back-end readout pixels as well as thescintillator elements used for PET and/or SPECT and/or Compton detectorscan be different from the dimensions of the readout pixels andscintillator elements used for CT (and tomosynthesis) detectors. (Notethat, as previously described herein, Compton detectors can be used asPET and/or SPECT detectors and vice versa. If only PET and/or SPECTdetectors are described then in at least one implementation they canincorporate Compton detector features).

The edge-on, n/m-level CT-PET, CT-SPECT or CT-SPECT-PET (including ringand cone beam CT, tomosynthesis) detector allows for many configurations(including Compton implementations of PET and/or SPECT). For example,replace one level (level n) of the n-level multispectral CT detectorwith one additional level of a PET scintillator (along with anyappropriate changes in the readout photodetector and electronics) ineither the extended edge-on detector array configuration or the n-leveledge-on detector array configuration followed by a 3D scintillatordetector stack). If the level n CT scintillator employed integratingreadout electronics with a photodiode then these items can be replacedwith a photodetector offering gain (such as SiPMs, silicon nanowires,APDs, etc.) and readout electronics capable of at least one of energyintegration or photon counting as well as photon counting with energyresolution. The collimated CT radiation is preferably incident on anentry side face whereas gamma radiation typically experiences limited orno collimation and can be incident on multiple side faces (including thephotodetector-coupled side face).

In general, the depth of the additional level of PET scintillator can beimplemented so as to provide acceptable detection efficiency (stoppingpower, fluorescence light collection) for the CT spectrum present atthat level and thus can be different from the old level n CTscintillator depth as well as any of the m levels of PET scintillatordepths. The limiting case involves replacing all n levels of CTscintillators (including the case n=1) with n levels of PETscintillators of appropriate depths. The PET scintillator(s) employedwith the m-levels of PET scintillators need not be the same as the PETscintillator(s) employed with the replaced level(s) of CT scintillators.

The implementation of one or more PET scintillator materials istypically based on various physical properties such as energy-dependentstopping power, fluorescence efficiency, fluorescence spectrum,transparency and index of refraction, temporal decay, radiationhardness, hydroscopic nature, etc. as well as cost The edge-on designwill typically include some internal collimation to reduce x-ray crosstalk and shield readout electronics (such as ASICs) from direct CT (ortomosynthesis) radiation. An external collimator (which can be mademobile if desired) is typically used for CT to provide x-ray scatterreduction although an alternative is to implement numerical algorithmsmay be employed to estimate and correct for x-ray scatter. In someimplementations PET detectors implement internal collimation. Thepresence of any internal as well as any external detector collimationshould be accounted for when acquiring simultaneous or sequential CT (ortomosynthesis) and PET images.

FIG. 19A illustrates an edge-on, n/m-level CT-PET detector 521 for n=1(photodetectors 286 coupled to the side face of scintillators 276) andm=1 in which the PET detector implements photodetectors 296 coupled tothe side face of scintillators 291.

FIG. 19B illustrates an edge-on, n/m-level CT-PET detector 521 for n=2(photodetectors 286, 287 coupled to the side faces of scintillators 276,277) and m=1 in which the PET detector implements photodetectors 296coupled to the side face of scintillators 291.

FIG. 19C illustrates an edge-on, n/m-level CT-PET detector 521 for n=3(photodetectors 286, 287, 288 coupled to the side face of scintillators276, 277, 278) and m=1 in which the PET detector implementsphotodetectors 296 coupled to the side face of scintillators 291.

FIG. 19D illustrates an edge-on, n/m-level CT-PET detector 521 for n=3(photodetectors 286, 287, 288 coupled to the side face of scintillators276, 277, 278) and m=2 in which the PET detector implementsphotodetectors 296 coupled to the side face of scintillators 291.

FIG. 19E illustrates an edge-on, n/m-level CT-PET detector 521 for n=2(photodetectors 286, 287 coupled to the side faces of scintillators 276,277) and m=6 in which the PET detector implements photodetectors 296coupled to the side face of scintillators 291.

FIG. 19F illustrates an edge-on, n/m-level CT-PET detector 521 for n=1up to 6 (photodetectors 291 coupled to the side faces of scintillators291) and m=6 in which the PET detector implements photodetectors 296coupled to the side face of scintillators 291. One or more PET detectorlayers are therefore employed as CT detector layers.

FIG. 19G illustrates (from an end perspective) a shared edge-on,n/m-level CT-PET detector 521 for n=2 (photodetectors 286, 287 coupledto the side faces of scintillators 276, 277 in aphotodetector-scintillator, photodetector-scintillator symmetricarrangement) and m=4 in which the PET detector implements photodetectors296 coupled to the side face of scintillators 291.

A gap can be implemented between the two adjacentphotodetector-scintillators. An alternative implementation employs thetwo adjacent edge-on, dual layer CT detector elements in aphotodetector-scintillator, scintillator-photodetector asymmetricarrangement (enabling different maintenance and cooling techniques). ThePET detector elements are positioned behind the two adjacent edge-on,dual layer CT detector elements (in contrast to FIG. 19E).

The CT and PET detector depths shown in FIGS. 19A-19G are forillustrative purposes and therefore are not necessarily to scale.Optionally, one or more PET detector layers can be employed as CTdetector layers in an edge-on, n/m-level CT-PET detector. Furthermore,one or more CT detector layers can be employed as PET detector layers inan edge-on, n/m-level CT-PET detector.

Sharing one or more PET and CT detector layer without modifications may,depending on the properties of the shared layers, resulting in poorerperformance for at least one of the CT and/or PET detectors. If deemedbeneficial detector element properties within a shared detector layer(such as scintillator dimensions, materials, coatings, photodetectors(and readout electronics) can be selected to provide improved detectorresults for CT and/or PET. For example, one or more parameters such as asmaller scintillator volume, a brighter scintillator phosphor, a moreefficient photodetector, configurable readout electronics (spectral orPC, integration), etc. could be adjusted in order to compensate for thelower photon energies used in x-ray CT versus PET.

Although FIGS. 19A-G show PET detector layers with uniform propertiesconcerning scintillators and photodetectors (PET detector elements) itshould be understood, in general, that these PET detector elements arenot constrained to possess uniform properties. The PET photodetectorscan be implemented as pixelated, strip, or area photodetectors dependingon cost and performance requirements. The edge-on n-level scintillatordetector can be replaced by an edge-on n-level semiconductor detector(including a structured mold semiconductor detector). Yet anotherimplementation replaces the edge-on m-level scintillator detector withan edge-on m-level semiconductor detector (including a structured moldsemiconductor detector).

Yet another implementation replaces both of the edge-on n-level andm-level scintillator detectors with edge-on n-level and m-levelsemiconductor detectors (including structured mold semiconductordetectors). Yet another implementation replaces at least one of theedge-on n-level or m-level scintillator detector layers with an edge-onor face-on semiconductor detector (including structured moldsemiconductor detector) layer. Optionally, at least one energy-dependentradiation attenuation filter can be included with thesescintillator-scintillator, scintillator-semiconductor andsemiconductor-semiconductor detector implementations. Structured moldsemiconductor detectors, as previously described, can incorporate one ormore energy-dependent radiation attenuation filters.

If cost is an issue the edge-on, n/m-level CT-PET detector can beimplemented with an m=1 PET detector in an edge-on geometry or an m=1PET detector in a face-on geometry. The m=1 PET detector in a face-ongeometry can implement DOI resolution by implementing photodetectors atboth ends of the scintillator rods. Alternatives that can provide DOIresolution with photodetectors at one end of the scintillator rodsinclude, but are not limited to, a discrete or continuous phoswichdesign, various light sharing designs between adjacent discrete orvirtual rods (including the, 3D scintillator designs described herein)or applying a pattern of WLS materials to the scintillator rod surfacesuch that wavelength and/or pulse shape properties vary with positionalong the lengths of the scintillator rods. A previous implementationused m=2 in which 2 discrete scintillator rods segments with differentpulse properties were optically-coupled to form individual phoswichscintillator rods with two level DOI resolution. If multi-energy CT isnot required then an n=1 CT detector in an edge-on geometry or an n=1 CTdetector in a face-on geometry can be implemented.

The design principles detailed herein of integrated CT-PET, CT-SPECT andCT-SPECT-PET (ring or cone beam CT, tomosynthesis) detector modules withsimultaneous or sequential acquisition can be readily extended toinclude integrated SPECT-PET detector modules with simultaneous orsequential acquisition. Full ring and partial ring geometries ofdetectors can be implemented for ring CT. Integrated SPECT-PET detectormodules can implement one or more levels of scintillator(s) with 2D or3D resolution suitable for use with one or more SPECT radionuclides(with lower energies than 511 keV) backed by stacked structured 3Dscintillator PET detectors (or other 2D/3D PET detectors). Variations ofSPECT-PET detector modules include implementations with one or morelevels of semiconductors (or structured semiconductors) or a combinationof scintillator and semiconductor (structured semiconductor) levels. Forexample, a scintillator level used for SPECT or SPECT and CT could bereplaced by a semiconductor level.

As described earlier herein, in one implementation detectors such asstacked structured 3D scintillators can incorporate both SPECT and PETscintillators (or only PET scintillators) within a stack since theextreme count rates associated with x-ray CT are not an issue. Anotherimplementation modifies the (extended) edge-on, n/m-level CT-PETdetector design taught earlier herein such that the n levels of edge-onSPECT scintillator(s) are appropriate for the one or more SPECTradionuclide energies backed by m levels of edge-on PET scintillator(s)(an edge-on, n/m-level SPECT-PET detector) as an alternative to stackedstructured 3D scintillator PET detectors (or other 2D/3D PET detectors).One or more scintillator levels can be replaced by one or moresemiconductor (or structured semiconductor) levels. While ring orpartial ring (or cylindrical/spherical) detector geometries are oftendesirable, alternative detector geometries can employ planarimplementations of the detectors described herein (e.g. positronemission mammography) and in some cases a ring, etc. detector geometrycan be simulated by moving the planar detectors or arrangements ofplanar detectors (e.g. cone beam CT, SPECT).

Integrated CT-SPECT-PET imaging detectors can implement a combination ofone or more levels of scintillator(s) suitable for appropriate CT x-rayspectrum(s) with one or more SPECT radionuclides (with lower energiesthan 511 keV) backed by stacked structured 3D scintillator PET detectors(or other 2D/3D PET detectors). One or more scintillator levels can bereplaced by one or more semiconductor levels. As described previouslyherein, in one implementation detectors such as stacked structured 3Dscintillators can incorporate both SPECT and PET scintillators (or onlyPET scintillators) within a stack since the extreme count ratesassociated with x-ray CT are not an issue.

Yet another implementation modifies the (extended) edge-on, n/m-levelCT-PET detector design taught earlier herein in which one or more of then levels of edge-on CT scintillator(s) are appropriate for the one ormore SPECT radionuclide energies backed by m levels of edge-on PETscintillator(s) detectors (an edge-on, n/m-level SPECT-PET detector) asan alternative to stacked structured 3D scintillator PET detectors (orother 2D/3D PET detectors). The choice of CT scintillators and SPECTscintillators (and pixel depths) may in some instances represent acompromise since the most common SPECT radionuclide (Tecnetium-99 140keV gamma ray) has an energy at or above the energy limit for typical CTbeam spectra. Higher SPECT radionuclide energies can be covered by PETscintillator detectors.

As taught earlier herein, at least one level (level n) of the n-levelmultispectral CT/SPECT or SPECT detector can be replaced with oneadditional level of a PET scintillator (along with any appropriatechanges in the readout photodetector and electronics) in either the(extended) edge-on detector array configuration or the n-level edge-ondetector array configuration followed by a 3D scintillator detectorstack. Shared edge-on, n/m-level SPECT-PET, CT-PET, CT-SPECT andCT-SPECT-PET detector configurations can be implemented.

Implementations of CT-SPECT-PET detector modules include designs inwhich SPECT and/or PET detector resolution are comparable to, or coarserthan CT detector resolution for patient imaging (or small animalimaging). SPECT and PET detectors can be shared (rather than only PETdetectors). The edge-on design can include internal collimation toshield photodetectors and/or electronics from direct CT radiation.Internal shielding can be implemented in order to limit radiation crosstalk. External collimation (which can be made mobile if desired) can beimplemented for CT imaging as well as SPECT imaging. In someimplementations PET detectors implement internal collimation. Thepresence of any internal as well as any external detector collimationshould be accounted for when acquiring simultaneous or sequentialSPECT-PET, CT-PET, CT-SPECT or CT-SPECT-PET images.

Energy-dependent MTF (Modulation Transfer Function) or MTF(E)calibration techniques have been previously described (Nelson, U.S. Pat.No. 6,583,420) for various x-ray detectors used in slit scan x-raymammography, tomosynthesis, area imaging and CT (ring, cone beam),whether the energy resolution is inherent to the detector (spectroscopiccapability), is the result of the detector geometry or both. When energyresolution is only the result of the detector geometry then one or moredetector elements function as energy integrators or photon counters andenergy resolution is inferred for those detector elements according tothe energy-dependent attenuation properties of detector elements and theenergy spectrum of the incident radiation field upon the detectorelements.

An example of a detector in which energy resolution is only the resultof the detector geometry include a scintillator-based, edge-on, n-levelCT detector operating in the energy integrating or photon counting mode.An example of a detector in which energy resolution is both inherent andthe result of detector geometry is an edge-on n-level CT (ortomosynthesis or slit/slot) semiconductor/structured semiconductordetector (with energy resolution) followed by at least one level of ascintillator (or semiconductor/structured semiconductor) detectoroperating in an energy integrating or photon counting mode.

MTF(E) calibration techniques can be applied for dedicated multi-energyCT (and tomosynthesis, slit/slot, area) detectors as well as integratedCT-PET, CT-SPECT or CT/SPECT/PET detectors. Calibration of the SPECTand/or PET detectors present in an integrated detector usingconventional or unconventional SPECT and/or PET calibration phantoms andsources (Nelson, U.S. Pat. No. 6,583,420) should include effects fromnearby detectors (including CT detectors) and any collimation that willbe present during SPECT and/or PET imaging. In addition, if anydetectors incorporate radioactive materials then their impact needs tobe accounted for if deemed significant.

Ionizing Radiation Detectors Employing Multiple Information Carriers

When a detector interacts with ionizing radiation such as x-rays, gammarays, charged particles, neutral particles one or more local electric,magnetic, optical, acoustic, thermal properties of the detector materialmay be altered. For example, phenomena such as x-ray inducedphotoacoustics, magnetoacoustics, etc. are well-documented in theliterature. Ionizing radiation detector encoding techniques (implementedwith appropriate readout sensors) for properties such as spatial and/ortemporal resolution are not limited to optical passive methods such aslight sharing, the use of WLSs, etc.

Ionizing radiation detectors can employ one or more active and/orpassive encoding techniques that utilize non-optical informationcarriers with or without passive optical encoding techniques. Passiveencoding techniques could include modifying at least one of local orglobal electric, magnetic, optical, acoustic and thermal ionizingradiation detector properties (including, but not limited to, theionizing radiation detectors described herein such as slab, block,pixelated, array and structured detectors). Passive and/or activeencoding techniques can be implemented with ionizing radiationstructured detectors. These techniques are not limited to radiationdetectors and can be applied to the interrogation of material samplesincluding human tissue (using at least one of ionizing or non-ionizingradiation detectors).

One or more properties of ionizing radiation structured detectormaterials (including nanoparticle properties if employed) and even thestructured detector frames (including features such as hole or trenchsize, shape, surface properties, coatings, anode and/or cathodeproperties (if present), the incorporation of embedded sensors) can bemodified or enhanced in order to support the use of more than one typeof information carrier. For example, in one implementation anode and/orcathode materials used in an ionizing radiation structured detectorcould incorporate both conductive and acoustic properties (enabling theuse of both charge and acoustic information carriers). In anotherimplementation a hole or trench anode (for example) could be segmentedalong its depth providing DOI (as well as energy and/or particlesensitivity as a function of depth) in a face-on geometry or SAR in anedge-on geometry. An alternative to employing a single semiconductor ornanoparticle material in a segmented hole or trench in a face-ongeometry is to enhance energy and/or particle detection sensitivity byfilling successive hole or trench segments with a specific semiconductoror nanoparticle material with favorable energy or particle interactioncapabilities for at least a part of the incident radiation spectrum. Arelated concept is employed in dual energy CT in which a low Z phosphortop layer preferentially detects low energy x-rays followed by a high Zphosphor bottom layer for detection of moderate-to-high energy x-rays.

An active encoding technique is to embed or couple individual and/orarrays of sensors (probes) within or on the surface(s) of an ionizingradiation detector (including, but not limited to, the ionizingradiation detectors (scintillator, semiconductor, structured, gas,superconductor) described herein with configurations such as slab,block, pixelated, array, structured, 3D structured, layered, etc.). Thesensors could be at least one of: electrically-sensitive,magnetically-sensitive, electromagnetically-sensitive,acoustically-sensitive, temperature-sensitive, particle-sensitive. Thesensors could provide at least one of: temporal information, spatialinformation, temperature information, energy deposition information,polarity information, particle identity information, detector healthinformation in addition to detector readout information.

For example, an embedded sensor array could implement at least one of asingle-sided readout format, a double-sided readout format, anautonomous readout format. A single-sided readout could have one or moresensor elements positioned at the end or along the length of acable/wire at known positions and thereby encode spatial resolutioninformation. One or more cable/wires could extend to select distanceswithin the detector volume or traverse the detector volume. Adouble-sided readout format could employ sensor cables/wires thattraverse the detector volume from one side of a detector to the otherside (in one version employing a signal processing technique such assignal division to determine the position of an ionizing event withrespect to an individual sensor cable/wire). 2D and 3D arrays ofcables/wires can be implemented.

Wired and/or wireless readout capabilities can be implemented. Signalstrength (energy resolution) and timing information (temporalresolution), as well as other signal event information, may also beavailable depending on the capabilities of the sensor cable/wire. Ifmultiple sensor cables/wires are present then readout information fromthe various sensor cables/wires can be combined or correlated,potentially enhancing spatial resolution, etc. and (in some instances)enabling tracking and/or a more detailed analysis of the event. Passiveand/or active encoding techniques can be employed with, but are notlimited to, the ionizing radiation detectors described herein.

An additional active encoding technique is to introduce one or morestatic or dynamic virtual structure within or on the surface of thedetector. Sources of static or dynamic virtual structures can be theresult of applied external magnetic fields, electric fields,electromagnetic fields, acoustic fields, and/or temperature variations.For example, optical source(s) and/or an acoustic source(s) can impose astatic or dynamic wave form pattern within an ionizing radiationdetector (a simple pattern is a repetition of pulses). When the patternis disrupted due to an ionizing event the change in the pattern can bedetected optically and/or acoustically, providing at least one ofspatial resolution, temporal resolution, energy resolution or otherproperties of the ionizing event.

Optical and/or acoustic elements or arrays that generate a wave formpattern can be coupled to one or more surfaces of an ionizing radiationdetector including, but not limited to, semiconductor or scintillatorionizing radiation detectors and used to reconstruct the location and/ortiming of an ionizing event from the disruption of the pattern (notethat an acoustic source may also be used as a receiver). If the patternis dynamic then information concerning parameters such as position ortiming may be encoded into the waveform imposed on the ionizingradiation detector and thus extracted when the dynamic pattern ismodified. The art of encoding propagating acoustic and electromagneticwave forms is also adaptable to the fields of acoustics, opticalcommunications, radar, seismology, etc.

Multiple active encoding techniques can be employed simultaneously withionizing radiation detectors. Furthermore, active encoding techniquesand passive encoding techniques (including passive techniques describedin this specification) can be employed simultaneously with ionizingradiation detectors. It should be understood that optical source anddetectors include not only visible but also near-infrared, infrared andterahertz sources and detectors.

Ionizing radiation detectors that incorporate active and/or passiveencoding techniques that employ at least two types of informationcarriers to detect ionizing radiation represent an alternative toconventional ionizing radiation detectors that typically use a singletype of information carrier (one exception is liquid Xenon which cangenerate measurable electronic signals and a fluorescence signals). Forexample, acoustic/ultrasound elements or arrays can be coupled to one ormore surfaces of an ionizing radiation detector including, but notlimited to, semiconductor or scintillator ionizing radiation detectorsand used to reconstruct the location of an ionizing event by listeningto acoustic sounds (the pressure signature) generated by thephotoacoustic event.

The types of ionizing radiation detectors include, but are not limitedto, the 1D, 2D, 3D, Compton camera, Compton-PET, CT-PET, CT, SPECT,SPECT-PET, CT-Compton-PET, etc. radiation detectors described herein.Applications include medical, industrial, homeland security andscientific imaging and/or analysis.

FIG. 20A illustrates a perspective of a discrete 3D scintillator andphotoacoustic PET detector module 531 irradiated with ionizing radiation109 face-on in which a scintillator block 350 is coupled to aphotodetector 250 on one surface and an acoustic array 440 is coupled toa different surface.

FIG. 20B illustrates a perspective of a discrete 3D scintillator andphotoacoustic PET detector module irradiated with ionizing radiation 109edge-on in which a scintillator block 350 is coupled to a photodetector250 on one surface and an acoustic array 440 is coupled to a differentsurface.

In yet another implementation the photodetector and acoustic array areboth coupled to the same surface of the scintillator block. FIG. 20Cillustrates a perspective of a discrete 3D scintillator andphotoacoustic PET detector module 531 irradiated with ionizing radiation109 face-on in which a scintillator block 350 is coupled to aphotodetector 250 and an acoustic array 440 at the same surface.

FIG. 20D illustrates a perspective of a discrete 3D scintillator andphotoacoustic PET detector module 531 irradiated with ionizing radiation109 edge-on in which a scintillator block 350 is coupled to aphotodetector 250 and an acoustic array 440 at the same surface.

Alternative implementations include reversing the order of thephotodetector array and acoustic array (if acceptable photodetectionefficiency is achieved) or employing a single readout device offeringboth acoustic and photodetection capability.

In FIG. 20A the single photodetector provides information concerning atleast one of energy resolution and/or temporal resolution whereas theacoustic array provides information concerning at least one of spatialresolution, temporal resolution, energy resolution. The use of a singlephotodetector may offer advantages in terms of overall cost and/orimproved performance. Spatial resolution can be modified by one or moretechniques including, but not limited to: selecting a different surfaceto couple the photodetector and/or acoustic array to, implementing anacoustic array element density that ranges from coarse to fine,employing one or more additional acoustic arrays coupled to additionalsurfaces, implementing a 1D or 2D photodetector, employing one or moreadditional photodetectors (including 1D or 2D photodetectors) coupled toadditional surfaces (in some configurations enabling SAR or DOImeasurements). Although scintillator blocks are depicted in FIGS.20A-20D, other scintillator geometries including, but not limited to,scintillator arrays can be implemented. At least one of area, strip, andpixel array photodetector geometries can also be employed. Thephotodetectors may be used to provide spatial resolution, andphotodetectors and/or acoustic arrays may be coupled to more than onesurface.

The use of acoustic/ultrasound elements or arrays is not limited toimplementing only receiver capabilities. Optionally, acoustic/ultrasoundtransmitter capabilities can also be implemented. Alternatively,acoustic/ultrasound elements or arrays could be integrated directly intothe ionizing radiation detector (consider the case of a silicon-basedradiation detector wherein the technology of silicon CMOS ultrasoundtransmitter and/or receiver chips is applied). Furthermore,acoustic/ultrasound transmitters that can be employed with ionizingradiation detectors can also be used with tissue (the sample) to createstationary or dynamic tissue density variations within a tissue volumethat can be evaluated by employing x-ray or gamma ray (or other ionizingradiation) phase imaging techniques using at least one ionizingradiation detector module (including, but not limited to, the ionizingradiation detector modules described herein). Conventional radiologicaland phase images can be acquired and compared. The ionizing radiationsource can be external and/or internal to the tissue). This method isreadily extended to include the analysis of samples composed ofmaterials other than tissue for non-medical imaging applications inscience, industry and inspection.

In general, one or more surfaces of a scintillator array, slab or blockcan be coupled to one or more acoustic/ultrasound elements or arrays (tomeasure acoustic information carriers) and one or more photodetectors(to measure optical information carriers). The acoustic/ultrasoundelements or arrays would provide information relevant for at least oneof spatial resolution, temporal resolution, energy resolution (whereinat least one of spectral resolution, PC, energy integration can beimplemented).

The photodetector(s), such as single photodetectors, position-sensitivephotodetectors and photodetector arrays, would provide at least one ofenergy resolution (alternatively, PC and/or energy integration can beimplemented), temporal resolution, spatial resolution. Thesephotodetectors can be positioned in previously herein-described detectorconfigurations including, but not limited to, edge-on, face-on, edge-onwith SAR, face-on with DOI, edge-on with encoding, face-on withencoding, discrete and semi-continuous structured 3D scintillators. Thisrepresents a flexible detector format since specific resolutioncapabilities for the acoustic/ultrasound elements or arrays or thephotodetectors can be emphasized or de-emphasized according to theperformance requirements of an application such as SPECT imaging, PETimaging, x-ray imaging, Compton imaging, particle detection, industrialimaging, homeland security.

In some instances resolution information from both acoustic and opticalinformation carriers can be combined beneficially. For example, ifspatial resolution information (even if limited) is available fromphotodetectors this information can, in some cases, augment spatialresolution information provided by the acoustic/ultrasound elements orarrays. If spatial resolution is largely or entirely determined by theacoustic/ultrasound elements or arrays then the number of photodetectorelements could be reduced, simplifying photodetector and readoutelectronics design and/or cost. Then the scintillator detector could beadapted for one or more features such as stopping power, energyresolution, temporal resolution either without the constraint ofoptically-determining spatial resolution or while allowing for reducedoptically-determining spatial resolution (the acoustic/ultrasoundelements or arrays are used to augment the spatial resolutioncapabilities of the photodetectors).

Non-imaging optics designs can be implemented in some cases. In somecases the use of fewer photodetectors or even a single photodetector canresult in lower photodetector (and readout electronics) costs and/or theuse of higher performance photodetectors (in some instances withsimplified features). For example, in some implementations 1D or 2DPMTs, micro-channel plates, SiPMs, etc. could be employed with muchlarger pixel sizes than are currently employed for use in small animaland human SPECT and/or PET imaging.

The use of one or more acoustic/ultrasound elements or arrays to providespatial resolution (and/or temporal resolution, energy resolution)information is not limited to scintillator detectors. Other ionizingradiation detector materials including, but not limited to,semiconductor detectors, super conductors, storage phosphors,nanoparticles, gases, liquids, etc. (and even traditional non-detectorsolid, liquid and gas materials) can be used with an acoustic/ultrasoundelements or arrays in order to provide (or augment) at least one ofspatial resolution, temporal resolution, energy resolution. The ionizingradiation detector may be tuned to enhance its response for a specificinformation carrier.

For example, the acoustic response and/or acoustic transport of anionizing radiation detector can be enhanced by various techniquesincluding the selection of detector materials with inherent (or that canbe engineered) favorable acoustic properties (acoustic resonances,channeling) or the implementation of structured detector with enhancedacoustic properties. These passive encoding techniques, as describedearly, can be combined with active encoding techniques. Traditionalnon-detector materials include, but are not limited to, metals,ceramics, glasses, various forms of carbon, plastics, water, ice, etc.which can all function as ionizing radiation detectors. Optionally,uniform or non-uniform temperature control of detector materials and/ornon-detector materials can be implemented to enhance photoacousticresponse. Imposing a non-uniform temperature distribution within thematerial volume could be used to spatially-modify the photoacousticresponse of the material.

Transformer PET

Various implementations of PET scanners (uniform and non-uniformdetectors; ring, ring with flat detectors, split-ring,dual-ring/open-ring, slant-ring, flat panel (or other) detectorgeometries; PET, CT-PET, SPECT-PET, CT-SPECT-PET, Compton-PET,CT-Compton-PET, Compton-PET-SPECT, CT-Compton-PET-SPECT) can beimplemented. For example, an open-ring PET scanner designed forradiation therapy with PET or CT with PET implements two identical,fixed PET detector rings that are used cooperatively (coincidences canbe detected within the individual ring detectors as well as between ringdetectors). There is a gap between ring detectors in order toaccommodate a radiation therapy or CT source and an opposing detector(and/or a beam stop). Sections of the gap that are not exposed to theradiation source can be filled-in with PET detector arrays.

The open-ring PET scanner design can lack flexibility in some respects.A more flexible PET system design employs two or more PET scanners(which may implement the same or different detector geometries and/ordetector properties) that can operate independently or cooperatively(referred to as transformer PET or T-PET). Gaps between adjacent PETscanners (if deemed undesirable) can be reduced by implementingremovable cowling from one or both sides of adjacent PET scanners.

For example, a T-PET system comprised of at least two ring PET scanners(for the case of two ring PET scanners one can be fixed and one can bemobile or both can be mobile) could be used simultaneously to image asingle region of a patient with greater detector efficiency or anextended region of a patient (an extended axial field of view). Anextended region of a patient may include sub-regions with different PETdetector requirements (such as a patient's chest/heart and head orneck/head). Two or more ring PET scanners can be employed in place of adedicated whole body PET scanner. The ring PET scanners can be the sameor they can be different (for example, different diameters and/ordifferent detector properties). None, one or more ring PET scanners canbe employed with one or more non-ring PET scanners (planar, squarecylinder, rectangular cylinder, etc. geometries) to form a T-PET system.

The at least two ring PET scanners of the T-PET system could operateindependently (with the same patient or different patients) orcooperatively with coincidence detection implemented between ring PETscanners. Cooperative coincidence event processing can be implementedelectronically (detected event information can be shared dynamicallybetween PET scanners) or through off-line processing using detectedevent information. Whether the at least two PET scanners operateindependently or cooperatively with the same patient, optional shieldingcan be positioned between the PET scanners to reduce their effectivefield of view of radiation sources within the imaged volume.

FIG. 21 illustrates a T-PET system 601 in which a heart/chest PETscanner 620 and a head/neck PET scanner 630 operate cooperatively tosimultaneously acquire PET cardiac and brain images of a patient 600.When the two ring PET scanners are operated independently with the samepatient then optional stationary and/or adaptive shielding can beemployed to help limit the field of view for each PET scanner.Furthermore, the two ring PET scanners can operate cooperatively whenonly one region of the patient is imaged and thereby continue to benefitfrom the inherent improvement of overall detection efficiency(potentially reducing patient imaging time and/or the amount ofradioactivity given to patient).

If three or more PET scanners are present all PET scanners can beoperated cooperatively or independently as well as various combinationsof cooperative and independent PET scanners. If one or more of the PETscanners is to be operated independently the PET scanner(s) canoptionally be shielded from the other PET scanners using stationaryand/or adaptive shielding (extendable/retractable) to help limit thefield of view for the independently-operated PET scanner(s).

The improved flexibility of the T-PET system allows for the (at least)two ring PET scanners to be employed to image one patient's chest/heartand another patient's brain (or neck/brain) independently. The two PETscanners could offer the same or different levels of performance. Forexample, the brain or neck/brain PET scanner could optionally implement:detectors offering higher spatial resolution (including non-uniformspatial resolution), different detector orientations, materials, energyresolution, temporal resolution; a smaller ring diameter, differentreadout electronics, etc. than the chest/cardiac PET scanner.

Implementations of the T-PET system are not limited to ring detectorgeometries (including partial ring geometries) and can be readilyimplemented with other PET detector geometries as previously describedherein (including, but not limited, to flat panel geometries).Furthermore, T-PET systems can implement multiple PET detectorgeometries and detector properties (for example, mixing ring and flatpanel (or other) detector geometries; uniform and non-uniform detectors;the implementation of at least one of CT-PET, SPECT-PET, CT-SPECT-PET inat least one of the PET scanners).

A limiting case for T-PET system employs two or more stationary PETscanners which can operate cooperatively or independently or acombination thereof. If one or more of the PET scanners is to beoperated independently that (those) PET scanner(s) can optionally beshielded from the other PET scanners using stationary and/or adaptiveshielding to help limit the field of view for the independently-operatedPET scanner(s). If SPECT capabilities are present in the T-PET systemthen various combinations of SPECT-PET imaging can be implemented: allscanners implement only PET, all scanners implement only SPECT, allscanners implement only SPECT-PET, at least one scanner implements PET,at least one scanner implements SPECT, at least one scanner implementsSPECT-PET. Collimation and shielding can be employed as needed. Theflexibility of the T-PET system allows for (at least) two SPECT-PETscanners to be employed to image (at least) two patient independentlyusing at least one of SPECT, PET or SPECT-PET functionality.

Gas-Based Compton Scatter Pet Detectors

The use of gas-based, multilayer straw detectors as Compton scatterdetectors for PET imaging was described earlier herein. An advantage isthat multilayer straw detectors are straightforward to implement usingexisting technology. The Compton scattering geometry for a face-on (oran edge-on) multilayer straw detector can be improved by requiring thatthe lower (bottom) surface wall of a layer of straws utilizes the upper(top) surface wall of the straws that define the next layer within astack and that side walls within a layer of straws are shared, forming amonolithic multilayer straw detector.

FIG. 22A illustrates a planar, face-on implementation of a monolithicmultilayer straw detector 900 comprised of straw fibers with squarecross sections and shared walls and incident ionizing photon radiation109. Wires 905 function as anodes. The walls 910 function as a Comptonscattering material and as a cathode. Event location along a wire can bedetermined by charge division or timing techniques.

The side walls of the straw fibers used in a monolithic multilayer strawdetector can be thinned or eliminated entirely Eliminating side wallssimplifies manufacturing but (as is the case of thinning side walls)reduces the total volume of low-Z material available for Comptonscattering. If the sidewalls within a layer are eliminated entirely theresult is a monolithic multilayer, multiwire proportional counterdetector with anode wires and cathode upper and lower surface wallsdefining each layer.

FIG. 22B illustrates a planar, face-on implementation of a monolithicmultilayer multiwire proportional counter detector 901 and incidentionizing photon radiation 109. Wires 906 function as anodes. The walls911 function as a Compton scattering material and as a cathode.Monolithic single layer multiwire proportional counter detectors canalso be implemented.

An alternative implementation (that employs much shorter anode wiresbetween crossed cathode strips) is a monolithic multilayer crossedstrips multiwire proportional counter detector. Yet anotherimplementation is to replace the array of wires in each layer with a gaselectron multiplier (GEM) foil or a micromegas mesh with a 2D anodeplane readout. Although more expensive than arrays of wires,improvements in spatial and timing resolution (and event ratecapability) are possible. Yet another implementation is a monolithicmultilayer microstrip gas chamber (including position-sensing microstripgas chamber variations).

Conventional RPCs, microstrip gas chambers, multiwire proportionalcounters and crossed strips multilayer proportional counters (includingreadout techniques to impose spatial resolution along the length or awire or strip), GEM and micromegas detectors are described. Monolithicsingle layer and multilayer straw, multilayer multiwire proportionalcounter (including crossed strips, GEM and micromegas variations) andmicrostrip gas chamber (including position-sensing microstrip gaschamber variations) detectors (including arrays) can be employed asstand-alone PET detectors or as front-end detectors along with one ormore layers of back-end detectors for PET imaging. Planar and/or curvedgeometries can be employed in edge-on and/or face-on orientation.

Monolithic single layer and multilayer straw detectors (as well asmultilayer multiwire proportional counter (including crossed strips, GEMand micromegas variations) and microstrip gas chamber (includingposition-sensing microstrip gas chamber variations) detectors) can beemployed as PET detectors in edge-on and/or face-on orientation. Forexample, a ring (or an approximation thereof) PET detector geometrycould be implemented using multiple monolithic single layer strawdetectors oriented edge-on along the axial direction or using multiplecurved monolithic multilayer straw detectors oriented face-on or edge-on(alternatively, multiple planar monolithic multilayer straw detectorsoriented in the axial direction could be distributed about thecircumference of the ring).

Typically, the straw, multiwire, microstrip, GEM, micromegas and RPC gasdetectors employed for Compton scattering offer at least one of:good-to-excellent spatial resolution and/or temporal resolution but withonly marginal or limited energy resolution for events originatingoutside the gas medium. Energy losses experienced by individualphotoelectric, Compton and secondary electrons that escape from thelow-Z scattering material into the gas are not predictable.

If excellent temporal resolution is available then TOF PET (offeringimproved image reconstruction resolution and reduced background noise)can be implemented on some level. Structured straw detectors (as well asother gas-filled detectors such as multiwire, microstrip and RPCdetectors) employing low-Z materials (glass, bakelite, plastic,Aluminum, etc.) are candidates for front-end detectors in extended axialfield of view whole body PET (with or without TOF capability) as well asPET (with or without TOF capability) since large area detectors can bemanufactured directly as a single detector piece or assembled frommultiple detector pieces.

The properties of the back-end detector can be customized to compensatefor any weakness of the structured straw, multiwire, microstrip, GEM,micromegas or RPC gas detector. For example, in one implementation, anRPC detector offers good-to-excellent spatial and temporal resolutionbut poor energy resolution. A back-end detector with adequate temporalresolution then should offer good-to-excellent energy resolution andspatial resolution.

For example, assume that tracking is implemented within and between thefront-end Compton scattering detector material (with poor energyresolution) and the back-end detector (with good energy resolution) andreasonable estimates for scattering angles (aside from Doppler effects)can be made, then the Compton equation can be applied one or more timesto approximate the incident energy of the gamma ray at the front-enddetector (which is relevant when an energy discrimination window isemployed). This method is particularly useful for the simplest case of asingle Compton-scattered photon event in the front-end detector thatterminate as a photoelectric event in the back-end detector since theaccuracy of any energy and spatial position measurements are limited.

Variants of these detector designs described herein can be furthercustomized for imaging charged and neutral particles with appropriatechanges to the front-end detectors and/or the back-end detectors (ifpresent). The use of these gas-based Compton scatter detectors is notlimited to PET imaging and they can be employed in any imagingapplication involving ionizing radiation including, but not limited to,nuclear medicine imaging (SPECT, PET-SPECT, Compton), medical x-rayimaging, homeland security imaging, industrial imaging, and scientificimaging.

High Resolution Detectors for Dental Imaging

Structured detectors such as structured mold detectors and transparentnanoparticle storage phosphor plate detectors can be employed in aface-on orientation for high spatial resolution (typically greater than7-10 lp/mm) applications such as intraoral dental x-ray imaging (seeNelson, U.S. Pat. No. 9,384,864 and U.S. patent application Ser. No.13/199,612, U.S. Publication No. 2012/0181437). Structured 3Dsemiconductor detectors (including amplified, co-doped and highresistivity variations) including, but not limited to, 3D Si, 3D GaAs,3D CdTe, 3D CZT, 3D Ge, 3D TlBr, 3D Se, 3D HgI₂, etc. (including dopedversions) that can be employed in various implementations of Compton,PET, SPECT, CT, slit, slot, and area x-ray imaging as described hereinare also suitable for applications including, but not limited to, highspatial resolution intraoral dental x-ray imaging and x-ray fluorescenceimaging. Structured 3D semiconductor detectors can also be referred toas 3D semiconductor detectors.

For intraoral dental imaging the structured 3D semiconductor detectorwould be employed in a face-on geometry. The same protective shell(frame) and electronic readout system described previously by Nelson(U.S. patent application Ser. No. 13/199,612, U.S. Publication No.2012/0181437) can be employed while the high resolution structured moldsemiconductor detector array is replaced with a high resolutionstructured 3D semiconductor detector array. The structured 3Dsemiconductor detector can be implemented in at least one ofintegrating, photon counting and spectral resolution detection modes.

The structured 3D semiconductor detector can offer advantages, in somecases, with respect to structured mold semiconductor detectors in termsof materials that can be utilized and manufacturability (includingmaterials with limited electron and/or hole mobility). Detectors ofreduced thickness compared to SPECT, PET or CT imaging can be employeddue to the relatively low x-ray energies used in dental intraoralimaging. The anode and cathode 3D structures created by drilling holesinto the semiconductor detector material allows for the creation of verysmall pixels in contrast to the manufacturing difficulty associated withstructured mold semiconductor detector which requires filling very smallholes with semiconductor or nanoparticle detector materials.

The components of a digital x-ray camera for intraoral dental imagingare shown in FIGS. 23A-23D.

FIG. 23A illustrates a perspective of a structured 3D semiconductorx-ray detector 921 with an array of holes 120 (anodes and cathodes) thatare electronically-coupled to an attached substrate incorporatingreadout circuitry 150 with a power and communication link 175 forconnection to a computer with incident radiation field 105.

FIG. 23B illustrates a perspective of a structured 3D semiconductorx-ray detector 921 with an array of segmented anode and cathode channels121 (rather than holes) that are electronically-coupled to an attachedsubstrate incorporating readout circuitry 150 with a power andcommunication link 175 for connection to a computer with incidentradiation field 105.

Channels can be continuous (spanning the length, width or diagonallength of the active detector area) or discontinuous (spanning a length,width or diagonal length less than that of the active detector area).The hole dimensions and channel dimensions as well as the number ofholes and channels shown are for illustrative purposes only. Bothchannels and holes can be employed within a structured 3D semiconductordetector.

FIG. 23C illustrates a perspective of a movable protective cover 127that can slide onto the protective shell or frame 126 holding astructured 3D semiconductor x-ray detector 921 with holes 120 andattached substrate incorporating readout circuitry as well as a powerand communication link 175 connected to a computer. Alternatively,protective shell or frame 126 can hold a structured 3D semiconductorx-ray detector with channels 121. Optionally, the power source andcommunication link can be contained within the protective cover (e.g. abattery, a wireless link).

FIG. 23D illustrates a perspective of a movable protective cover 127 inplace over the protective frame 126 forming a digital x-ray detector orcamera 941 for intraoral dental imaging with a power and communicationlink 175 connected to computer 176.

High resolution transparent nanoparticle storage phosphor sheetsdetectors (mounted on a base or plate) offer advantages compared toscintillator on CMOS/CCD detectors for intraoral dental imaging (greaterx-ray stopping power, greater spatial resolution). However, the densityof nanoparticles embedded within a transparent medium such as a glassceramic is often low. This requires thicker storage phosphor sheets(approximately 400-1,000 microns or more depending on desired x-raystopping power) to improve x-ray stopping power which in turn requiringa thicker protective shell (frame) that increases patient discomfort).

Yields often decrease as transparent nanoparticle storage phosphordetector sheet thickness increases. Uniform dispersion of nanoparticlestorage phosphors within the transparent medium can be problematic andthe nanoparticle storage phosphors make the transparent medium brittlewhich in turn requires a thicker protective shell (frame) or the use ofexpensive rigid materials in the shell (frame) to increase its rigidity.Furthermore, fluorescence efficiency of nanoparticle storage phosphorsmay typically be substantially less than for a more conventional storagephosphor, often potentially increasing patient radiation doseunnecessarily.

An alternative to the nanoparticle storage phosphor sheet detector thatdecreases detector sheet thickness (thickness of approximately 50-200microns are suitable for many intraoral dental imaging applications)while restoring fluorescence efficiency is to employ a conventionalstorage phosphor deposited as a transparent, continuous thin film orlayer (which can be applied with good uniformity) on a rigid plate orsupport or glued to a rigid plate or support (a substrate). The rigidplate or support can be implemented as a removable structure from theprotective shell (frame) or as an integral part of the protective shell(frame) which then requires a moveable protective shell lid or cover forscanning and readout. The rigid plate surface adjacent to the storagephosphor layer can implement at least one of reflectors, absorbers,scatterers, and WLS materials.

The same protective shell (frame), nanoparticle storage phosphordetector scanning mechanism and electronic readout system can beemployed as well as modifications (etchings, coatings, etc.) to one orboth of the transparent storage phosphor film surfaces and/or thestorage phosphor film rigid plate or support (previously describedherein for one or both surfaces of the nanoparticle storage phosphorplates as well as the supporting surface). The protective shell (frame)thickness can be kept the same or decreased (which improves patientcomfort) due to the reduced thickness of the storage phosphor filmcompared to the nanoparticle storage phosphor plate.

A dual energy implementation of the storage phosphor film detectoremploys transparent storage phosphor film detectors (of the same storagephosphor material or different storage phosphor materials) mounted onboth sides of support plate. If two different storage phosphor materialsare employed the material with the lower effective Z is typicallypositioned to intercept the x-ray beam first. If only one storagephosphor material is employed the layer that intercepts the x-ray beamfirst is typically of the same or reduced thickness relative to thesecond layer. The two surfaces can be scanned and read out at the sametime (e.g., using two independent readout systems), or sequentially.Optionally, the support plate can provide energy filtration by selectiveadsorption of the incident radiation, so that the energy spectrum ismodulated on passage through the support plate. An alternativeimplementation (which may limit the choice of storage phosphors) employsa stack of at least two different transparent storage phosphor films(different stimulating wavelengths and/or different emission spectrums),minimizing registration issues and reducing the thickness of theprotective frame. Multi-energy implementations employ at least oneadditional transparent storage phosphor film layer on at least one sideof the support plate. Furthermore, variations of dual energy (ormulti-energy) implementations include, but are not limited to, replacinga continuous transparent storage phosphor film layer with either agranular storage phosphor layer or a transparent nano-particle storagephosphor layer or a pixelated storage phosphor layer or a structuredstorage phosphor layer.

FIG. 24A illustrates a perspective of a flat, small area transparentstorage phosphor film 10 on a support plate 12 (a transparent storagephosphor film support plate) that can be mounted within a protectiveframe to provide a detector system 951.

FIG. 24B illustrates a perspective (transparent edge-on) of atransparent storage phosphor film support plate 132 attached to aprotective frame 131 with a movable protective layer 135 that slidesonto the protective frame to provide a detector system 961.

FIG. 24C illustrates a perspective of a protective frame with a base165, sidewalls 164 and a rotatable movable protective layer or cover 163attached by a hinge 167 to the protective frame. A transparent storagephosphor film support plate can be made separable from the frame or itcan be integrated into the frame to provide a detector system 971.

EXAMPLES

Suitable examples and embodiments of the invention include, but are notlimited to, any one or more of the following. Individual features ofthese examples and embodiments can be provided in any order orcombination suitable to radiation imaging, as described herein.

A detector module comprising: a first layer of scintillator rods, eachscintillator rod in the first layer extending in a first direction; anda second layer of scintillator rods stacked on or adjacent the firstlayer, each scintillator rod in the second layer extending in a seconddirection; wherein the first and second directions are transverselyoriented, such that the first and second layers of scintillator rods arecrossed to define a region of light sharing between the scintillatorrods of the first layer and the scintillator rods of the second layer.

In particular examples and embodiments, the scintillator rods in thefirst layer have at least one of different dimension or differentscintillator materials with respect to the scintillator rods in thesecond layer, such that the first and second layers have one or more ofdifferent spatial, timing and energy resolution. Scintillator rodswithin at least one of the first and second layers have at least one ofdifferent dimensions or different scintillator materials. At least someof the scintillator rods have different cross-sectional area transverseto at least one of the first and second directions, respectively. Atleast some of the scintillator rods have different longitudinaldimension along at least one of the first and second directions,respectively.

The detector module, further comprising encoding features disposed atleast one of in, on and between one or both of the scintillator rods ofthe first layer and the scintillator rods of the second layer, whereinthe encoding features are configured to modulate propagation of opticalsignals along one or both of the layers of scintillator rods, or betweenthe layers of scintillator rods. In various examples and embodiments,the scintillator rods are configured to generate the optical signals inresponse to one or more of x-ray radiation, gamma radiation, or particleradiation.

The detector module, further comprising a pattern applied to one or bothof the first and second layers of scintillator rods, or between thescintillator rods, the pattern comprising at least one of a reflective,diffusive, absorptive, WLS, photonic crystal, nano-layered metamaterials(including nanocavities), refracting, diffracting and lens opticalsurface adapted to modulate propagation of optical signals along one orboth of the layers of scintillator rods, or between the layers ofscintillator rods. In various examples and embodiments, the patterncomprises a grid of optical features having a spacing corresponding tothat of the crossed scintillator rods in the first and second layers,the grid of optical features configured to modulate lateral light flowtherebetween.

The detector module, further comprising an offset portion of one or bothof the layers of scintillator rods with respect to the other, whereinthe offset portion is defined by an extension of the respectivescintillator rods beyond a periphery of the other layer, outside theregion of light sharing between the first and second layer.

The detector module, further comprising a photodetector element coupledto the offset portion of one of both of the layers, wherein thephotodetector element is configured to sense optical signals from theextension of the respective scintillator rods, outside the region oflight sharing between the first and second layer. In various examplesand embodiments, the scintillator rods in the first layer and thescintillator rods in the second layer have at least one of asubstantially same cross sectional dimension, a substantially samelongitudinal dimension, and a substantially same scintillator materialwith a substantially same response.

A detector module comprising: a layer of scintillator rods, eachscintillator rod extending in a longitudinal direction therein; a pixelstructured scintillator layer stacked on or adjacent the layer ofscintillator rods, the pixel structured scintillator layer extendingtransversely to the longitudinal direction; and a light sharing regiondefined between the layer of scintillator rods and the pixel structuredscintillator layer.

A detector module comprising: a layer of scintillator rods, eachscintillator rod extending in a longitudinal direction therein; acontinuous scintillator layer stacked on or adjacent the layer ofscintillator rods, the continuous scintillator layer extendingtransversely to the longitudinal direction; and a light sharing regiondefined between the layer of scintillator rods and the continuousscintillator layer.

A detector module comprising: at least first and second layers ofcrossed scintillator rods configured to generate optical signals inresponse to ionizing radiation, the crossed scintillator rods extendingin first and second transverse directions in the first and secondlayers, respectively; a light sharing region defined between the firstand second layers, wherein the optical signals are transmitted betweenthe respective crossed scintillator rods; and a plurality ofphotodetector elements configured to convert the optical signals intooutput characterizing the radiation.

In various examples and embodiments, the first and second layers ofscintillator rods are defined in discrete, stacked scintillator elementshaving the light sharing region defined therebetween. The first andsecond layers of scintillator rods are defined in a unitary scintillatorelement having the light sharing region defined therein.

The detector module, further comprising a plurality of dividersextending in the first and second directions in the first and secondlayers respectively, the dividers defining the scintillator rods in therespective layers. In various examples and embodiments, dividers extendinto the unitary scintillator element from a first major surface of thefirst layer and from a second major surface of the second layer,respectively, each of the dividers having a depth less than a thicknessof the unitary scintillator element between the first and second majorsurfaces. The depth of the dividers varies between differentscintillator rods in one or both of the first and second layers. Thedividers comprise at least one of physical or virtual gaps formed in theunitary scintillator element, between the respective scintillator rods.The dividers comprise at least one of a reflective, diffusive,absorptive or WLS optical feature or material disposed between therespective scintillator rods.

The detector module, further comprising an encoding pattern configuredto modulate the transmission of the optical signals between the firstand second layers of scintillator rods. In various examples andembodiments, the photodetector elements are disposed at ends of thescintillator rods, the ends defining an end face having a crosssectional area transverse to the first and second direction in the firstand second layers, respectively.

A detector module comprising: at least first and second layerscomprising scintillator rods and a pixel structured scintillator,respectively, configured to generate optical signals in response toionizing radiation, wherein the scintillator rods extend in alongitudinal direction and the pixel structured scintillator extendstransversely thereto; a light sharing region defined between the firstand second layers, wherein the optical signals are transmitted betweenthe scintillator rods and the pixel structured scintillator; and aplurality of photodetector elements configured to convert the opticalsignals into output characterizing the radiation.

In various examples and embodiments, the first layer comprises aplurality of discrete scintillator rod elements coupled to the secondlayer; and the second layer comprises a plurality of discrete pixelstructured scintillator elements coupled to the first layer.

In various examples and embodiments, the first and second layers areformed of a unitary scintillator element having a first major surface ofthe first layer and a second major surface of the second layer, inwhich: the first layer comprises a plurality of virtual rod elementsdefined by a first set of dividers extending into the unitaryscintillator element from the first major surface thereof; the secondlayer comprises a plurality of virtual pixel structured scintillatorelements defined by a second set of dividers extending into the unitaryscintillator element from the second major surface thereof; and each ofthe first and second sets of dividers has a depth less than a thicknessof the unitary scintillator element between the first and second majorsurfaces.

The detector module of claim 25, wherein: the first layer comprises aunitary scintillator element coupled to the second layer, the unitaryscintillator element having a set of dividers defining the scintillatorrods as a plurality of virtual rod elements therein; and the secondlayer comprises of a plurality of discrete pixel structured scintillatorelements coupled to the first layer. In various examples andembodiments, the first layer comprises a plurality of discretescintillator rod elements coupled to the second layer; and the secondlayer comprises a unitary scintillator element coupled to the firstlayer, the unitary scintillator element having a set of dividersdefining the pixel structured scintillator as a plurality of virtualpixel elements therein.

A detector module comprising: at least a first layer of scintillatorrods or a pixel structured scintillator coupled to a second layer ofsemi-continuous structured scintillator sheet, each of the layersextending in first and second transverse directions and configured togenerate optical signals in response to ionizing radiation; a lightsharing region defined between the first and second layers, wherein theoptical signals are transmitted therebetween; and a plurality ofphotodetector elements configured to convert the optical signals intooutput characterizing the radiation.

A detector module comprising: a first layer of scintillator defining aplurality of scintillator elements, each of the scintillator elementsextending in a longitudinal direction along the first layer; a secondlayer of scintillator defining a substantially continuous scintillatorelement extending adjacent the first layer defining the plurality ofscintillator elements; a plurality of photodetector elements configuredto convert optical signals generated by the scintillator into outputcharacterizing radiation interacting in one or both of the first andsecond layers.

In various examples and embodiments, the plurality of scintillatorelements are defined by dividers extending into the scintillator from afirst major surface of the first layer, wherein a depth of the dividersdefines a thickness of the first layer. Dividers are formed by at leastone of a reflective, absorptive or WLS optical feature or materialdisposed between the respective scintillator rods. The depth of thedividers is nonuniform and defines a corresponding nonuniform thicknessof the first layer. The depth of the dividers is selected to define acurved or arcuate boundary between the first and second layers.

Any such detector module, further comprising an encoding patternconfigured to modulate light sharing between the first and second layersof the scintillator.

A scintillator detector module comprising: a first layer of elongatescintillator elements, each extending in a first longitudinal directionalong the first layer; a second layer of elongate scintillator elements,each extending in a second longitudinal direction along the secondlayer; and a plurality of photodetectors configured to generate outputcharacterizing optical signals generated by the scintillator elements inresponse to radiation interacting in one or both of the first and secondlayers.

A scintillator detector module comprising: a first layer of elongatescintillator elements, each extending in a longitudinal direction alongthe first layer; a second layer of pixel structured scintillatorelements extending in first and second transverse directions along thesecond layer; and a plurality of photodetectors configured to generateoutput characterizing optical signals generated by the scintillatorelements in response to radiation interacting in one or both of thefirst and second layers.

A detector module comprising: a first layer comprising a plurality ofone or both of: elongate scintillator elements extending in alongitudinal direction along the first layer, and pixel structuredscintillator elements extending in first and second transversedirections along the first layer; a plurality of photodetectorsconfigured to generate output characterizing optical signals generatedby the scintillator elements in response to radiation interactingtherein; a second layer comprising one or more semiconductor detectorelements oriented edge on or face on with respect to the radiation.

A detector module, further comprising an intermediate layer ofscintillator disposed between the first and second layers. In variousexamples and embodiments, one, two or more of the first, second andintermediate layers are formed of a unitary scintillator material andfurther comprising a plurality of dividers formed in the unitaryscintillator material, the dividers defining the scintillator elementstherein. One, two or more of the first, second and intermediate layersare formed of discrete scintillator materials joined together to formthe detector module. The intermediate layer comprises a plurality ofelongate scintillator elements. The elongate scintillator elements inthe intermediate layer are disposed at a skew angle with respect to theelements of the first or second layer. The intermediate layer comprisesa plurality of pixel structured scintillator elements. The intermediatelayer comprises a continuous scintillator layer.

The detector module, further comprising an intermediate layer disposedbetween the first and second layers, wherein the intermediate layercomprises a non-scintillator material. In various examples andembodiment, the intermediate layer is configured to modulate lightsharing between the scintillator elements in the first and secondlayers. The elements in the second layer are oriented transverse to theelements in the first layer such that the elements in the first andsecond layers are crossed. The elements in the second layer are orientedalong the elements in the first layer such that the elements in thefirst and second layers are parallel.

Any such detector module, further comprising wavelength shiftingmaterial disposed at least one of in, on and between the scintillatorelements and the photodetector elements, wherein the wavelength shiftingmaterial is adapted to shift an emission spectrum of the scintillatorelements to match a spectral response of the photodetector elements. Invarious examples and embodiments, the wavelength shifting material isadapted to shift the emission spectrum of Cherenkov light emitted by thescintillator elements in response to the radiation to match the spectralresponse of the photodetector elements. The wavelength shifting materialis adapted to reduce light trapping of a fluorescence signal generatedby the scintillator elements in response to the radiation. Thewavelength shifting material is adapted to modulate the optical signalsas a function of position along the scintillator elements.

The detector module, further comprising an optical coating appliedbetween the wavelength shifting material and a surface of one or more ofthe scintillator elements, the optical coating having an index ofrefraction selected to enhance internal reflection at the surface. Invarious examples and embodiments, the photodetector elements aredisposed to detect the optical signals emitted at a single side or endor at opposite sides or ends of one or more of the scintillatorelements. The photodetector elements are disposed to detect opticalsignals emitted at sides or ends of the scintillator elements and havetiming resolution adapted to distinguish signals reflected from oppositesides or ends thereof.

Any such detector module, wherein the layers are curved or generallyarcuate to provide a focused geometry of the scintillator module withrespect to a radiation source. In various examples and embodiments, thelayers define one or more generally cylindrical shell segments orientedalong an axial direction transverse to a direction of the radiationsource. The layers define one or more generally spherical shell segmentsdisposed generally transverse to a direction of the radiation source.

A detector module comprising: a first layer of scintillator elements,each extending in a first direction along the first layer; a secondlayer of scintillator elements, each extending in a second directionalong the second layer, wherein the second direction is orientedtransverse to the first direction such that the scintillator elements inthe first and second layers are crossed; an intermediate layer disposedbetween the first and second layers; and a plurality of photodetectorsconfigured to convert optical signals generated by the scintillatorelements into output characterizing radiation interacting in at leastone of the first, second and intermediate layers.

A detector module comprising: a first layer of scintillator elements,each extending in a first direction along the first layer; a pluralityof photodetectors configured to convert optical signals generated by thescintillator elements into output characterizing radiation interactingtherein; a second layer of one or more semiconductor detector elementsoriented edge on or face on with respect to the radiation; and anintermediate layer between the first and second layers.

In various examples and embodiments, the intermediate layer comprises aplurality of scintillator blocks positioned between and opticallycoupled to the first and second layers of crossed scintillator elements.The optical signals are transmitted through the intermediate layerbetween the first layer and the second layer. The intermediate layer isencoded to modulate transmission of the optical signals as a function ofposition along the scintillator elements. The first, second andintermediate layers are formed of a unitary scintillator element. Theintermediate layer has internal structure. The intermediate layer isformed of a different scintillator material from the scintillatorelements of one or both of the first and second layers. The intermediatelayer is substantially transparent. The intermediate layer is configuredto transmit the optical signals from a scintillator element in the firstlayer to multiple scintillator elements in the second layer. Thephotodetector elements are configured for the output to characterizeradiation of different energy spectra interacting in the first andsecond layers. The thicknesses of the first and second layers areselected to compensate for beam hardening.

Any such detector module, wherein one or more layers of the scintillatorelements comprise scintillator fibers or optical fibers. In variousexamples and embodiments, at least one of the layers comprisesscintillator material having optical fibers embedded therein. At leastone of the layers comprises a plurality of scintillator fibers oroptical fibers coupled to a scintillator sheet or block. At least someof the fibers comprise encoded cores. At least some of the fiberscomprise wavelength shifting materials provided in an encoded pattern.

Any such detector module, wherein event localization information isemployed to correct for optical signal propagation time within thedetector module utilizing at least one of a direct event signal, areflected signal, a cross-coupled signal, a wavelength shifted signal,and an indirect signal responsive to the event localization in thedetector module.

An edge-on, multispectral CT scintillator detector system comprising aplurality of such detector modules, wherein the radiation is incident onend faces of the scintillator elements in at least one of the layers.

In various examples and embodiments, the photodetector elements areoptically coupled to side faces of the scintillator elements in the atleast one of the layers. The first and second layers are responsive todifferent energy ranges of the radiation. The first and second layershave different energy responses. The first and second layers have a sameenergy response.

The detector system, further comprising at least a third layer ofscintillator elements having a different energy response from at leastone of the first and second layers. The detector system, furthercomprising at least a third layer of scintillator elements having a sameenergy response as at least one of the first and second layers. Invarious examples and embodiments, the radiation comprises a combinationof x-ray radiation from an x-ray source and gamma radiation from a gammasource.

The detector system, further comprising readout electronics coupled tothe photodetector elements, wherein the readout electronics are adaptedfor a combination of CT and at least one of PET, SPECT, PET-SPECT andCompton imaging. In various examples and embodiments, the radiationcomprises x-ray radiation from plural x-ray sources. Scintillatorelements in the first and second layers comprise different, relativelylower-Z and relatively higher-Z scintillator materials, respectively.

The detector system, further comprising a collimator disposed withrespect to the first and second layers, wherein the collimator isconfigured to modulate scattering of the radiation. The detector system,further comprising readout electronics coupled to the photodetectorelements, wherein the readout electronic are configured for one or moreof energy integration, photon counting, and photon counting with energyresolution. In various examples and embodiments, the detector system isconfigured for at least one of ring CT, cone beam CT, and tomosynthesisimaging.

An edge-on, integrated PET-CT detector system comprising a plurality ofany such detector modules, wherein the radiation is incident on endfaces of the scintillator elements in at least one of the layers and thephotodetector elements are optically coupled to side faces thereof. Invarious examples and embodiments, the radiation comprises a combinationof x-ray radiation from an x-ray source and gamma radiation from a gammasource. The first and second layers have different energy response tothe radiation. The first and second layers have a same energy responseto the radiation.

The detector system, further comprising at least a third layer ofscintillator elements having a different energy response to theradiation than at least one of the first and second layers. The detectorsystem, further comprising at least a third layer of scintillatorelements having a same energy response to the radiation as at least oneof the first and second layers. The detector system, further comprisingreadout electronics configured for determination of an interactionposition of the radiation within the scintillator elements.

In various examples and embodiments, scintillator elements in differentlayers have different cross sectional geometry. The detector modulesdefine a full ring geometry with respect to a source of the radiation.The detector modules define a partial ring geometry with respect to asource of the radiation. The detector modules define a planar geometry.

Any such detector system, further comprising an encoding patternconfigured to modulate light sharing between the first and secondlayers, wherein the encoding pattern is adapted to modulate the opticalsignals as a function of position along the first and second layers. Anysuch detector system, further comprising a pattern of wavelengthshifting material applied to the scintillator elements, wherein thepattern of wavelength shifting material is adapted to modulate theoptical signals as a function of position along the first and secondlayers. Any such detector system, further comprising readout electronicscoupled to the photodetector elements, wherein the readout electronicsare adapted for a combination of at least two of CT, PET, SPECT andCompton imaging.

A radiation detector module comprising: a scintillator elementconfigured to generate optical signals in response to incidentradiation; a photodetector coupled to at least a first surface of thescintillator element, the photodetector configured to convert theoptical signals into output characterizing the radiation; and anacoustic array coupled to at least a second surface of the scintillatorelement, the acoustic array configured to convert acoustic signalsgenerated in the scintillator element into output characterizingacoustic energy deposited therein.

In various examples and embodiments, the radiation is incident faceon-to the scintillator element, along a major surface thereof. Theradiation is incident edge-on to the scintillator element, along an edgedefined transverse to first and second major opposing surfaces thereof.The photodetector and the acoustic array are disposed on opposingsurfaces of the scintillator element. The photodetector and the acousticarray are disposed on a same surface of the scintillator element. Thephotodetector and the acoustic array are disposed on non-parallelsurfaces of the scintillator element. At least one of the photodetectorand the acoustic array is coupled to more than one surface of thescintillator element.

The detector module, further comprising readout electronics coupled tothe photodetector, the readout electronics configured to characterizethe incident radiation based on one or more of energy resolution,spatial resolution or temporal resolution of the optical signals. Thedetector module, further comprising readout electronics coupled to theacoustic array, wherein the readout electronics are configured tocharacterize the incident radiation based on one or more of spatialresolution, temporal resolution or energy resolution of the acousticsignals. In various examples and embodiments, the acoustic arraycomprises an acoustic transceiver configured for generating audio orultrasound signals.

An imaging system comprising any such radiation detector module adaptedto image at least one of an external or internal radiation field, andfurther comprising: a sample acoustically coupled to the detectormodule; an acoustic transceiver configured to generate densityvariations in the sample; and a processor in data communication with theradiation detector module, the processor configured to evaluate thesample based on the first and second outputs. In various examples andembodiments, the radiation detector module is configured for the imagingsystem to acquire conventional radiological images of the sample basedon the output characterizing the radiation and phase images of thesample based further on the output characterizing the acoustic energy.

A radiation detector module comprising: a detector element configured togenerate electronic and acoustic signals in response to incidentradiation passing through a sample; and an acoustic array or transceivercoupled to the detector element and acoustically coupled to the sample,wherein the acoustic array or transceiver is configured to convert theacoustic signals into output characterizing the incident radiation. Invarious examples and embodiments, the detector element comprises a solidstate detector medium. In various examples and embodiments, the detectorelement comprises a scintillator.

INCORPORATED REFERENCES

These references are expressly incorporated by reference herein:

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While this invention has been described with reference to exemplaryembodiments, it will be understood by those skilled in the art thatvarious changes can be made and equivalents may be substituted withoutdeparting from the spirit and scope thereof. Modifications may also bemade to adapt the teachings of the invention to particular problems,technologies, materials, applications and materials, without departingfrom the essential scope thereof. The invention is not limited to theparticular examples that are disclosed herein, but encompasses allembodiments falling within the scope of the appended claims.

The invention is thus susceptible to various modifications andalternative forms, specific examples thereof having been shown by way ofexample in the drawings and described in detail. It is understood thatthe invention is not limited to the particular forms or methodsdisclosed, but to the contrary, the invention encompasses allmodifications, equivalents, and alternatives falling within the spiritand scope of the claims.

1. A detector module comprising: a first layer of scintillator rods,each scintillator rod in the first layer extending in a first direction;and a second layer of scintillator rods stacked adjacent the firstlayer, each scintillator rod in the second layer extending in a seconddirection; wherein the first and second directions are transverselyoriented, such that the first and second layers of scintillator rods arecrossed to define a region of light sharing between the scintillatorrods of the first layer and the scintillator rods of the second layer;and wherein the scintillator rods are configured to generate the opticalsignals in response to one or more of x-ray radiation, gamma radiation,or particle radiation.
 2. The detector module of claim 1, wherein thescintillator rods in the first layer have at least one of differentdimension or different scintillator materials with respect to thescintillator rods in the second layer, such that the first and secondlayers have one or more of different spatial, timing and energyresolution.
 3. The detector module of claim 1, wherein scintillator rodswithin at least one of the first and second layers have at least one ofdifferent dimensions or different scintillator materials fromscintillator rods within another of the first and second layers.
 4. Thedetector module of claim 1, wherein at least some of the scintillatorrods have different cross-sectional area transverse to at least one ofthe first and second directions, respectively; or wherein at least someof the scintillator rods have different longitudinal dimension along atleast one of the first and second directions, respectively.
 5. Thedetector module of claim 1, further comprising encoding featuresdisposed in, on or between one or both of the scintillator rods of thefirst layer and the scintillator rods of the second layer, wherein theencoding features are configured to modulate propagation of opticalsignals at least one of along one or both of the layers of scintillatorrods and between the layers of scintillator rods.
 6. The detector moduleof claim 1, further comprising a pattern applied to one or both of thefirst and second layers of scintillator rods, or between thescintillator rods, the pattern comprising at least one of a reflective,diffusive, absorptive, WLS, or photonic crystal material, nano-layeredmetamaterials (including nanocavities), or a refracting, diffracting orlens optical surface adapted to modulate propagation of optical signalsalong one or both of the layers of scintillator rods, or between thelayers of scintillator rods.
 7. The detector module of claim 6, whereinthe pattern comprises a grid of optical features having a spacingcorresponding to that of the crossed scintillator rods in the first andsecond layers, the grid of optical features configured to modulatelateral light flow therebetween.
 8. The detector module of claim 1,further comprising an offset portion of one or both of the layers ofscintillator rods with respect to the other, wherein the offset portionis defined by an extension of the respective scintillator rods beyond aperiphery of the other layer, outside the region of light sharingbetween the first and second layer.
 9. The detector module of claim 8,further comprising a photodetector element coupled to the offset portionof one of both of the layers, wherein the photodetector element isconfigured to sense optical signals from the extension of the respectivescintillator rods, outside the region of light sharing between the firstand second layer.
 10. The detector module of claim 8, wherein thescintillator rods in the first layer and the scintillator rods in thesecond layer have a substantially same cross sectional dimension.
 11. Adetector module comprising: at least first and second layers of crossedscintillator rods configured to generate optical signals in response toionizing radiation, the crossed scintillator rods extending in first andsecond transverse directions in the first and second layers,respectively; a light sharing region defined between the first andsecond layers, wherein the optical signals are transmitted between therespective crossed scintillator rods; a plurality of photodetectorelements configured to convert the optical signals into outputcharacterizing the radiation; and an encoding pattern configured tomodulate the transmission of the optical signals at least one of alongone or both layers of scintillator rods and between the first and secondlayers of scintillator rods.
 12. The detector module of claim 11,wherein the first and second layers of scintillator rods are defined indiscrete, stacked scintillator elements having the light sharing regiondefined therebetween.
 13. The detector module of claim 11, wherein thefirst and second layers of scintillator rods are defined in a unitaryscintillator element having the light sharing region defined therein.14. The detector module of claim 13, further comprising a plurality ofdividers extending in the first and second directions in the first andsecond layers respectively, the dividers defining the scintillator rodsin the respective layers.
 15. The detector module of claim 14, whereinthe dividers extend into the unitary scintillator element from a firstmajor surface of the first layer and from a second major surface of thesecond layer, respectively, each of the dividers having a depth lessthan a thickness of the unitary scintillator element between the firstand second major surfaces.
 16. The detector module of claim 15, whereinthe depth of the dividers varies between different scintillator rods inone or both of the first and second layers.
 17. The detector module ofclaim 15, wherein the dividers comprise at least one of a reflective,diffusive, absorptive, refractive, diffractive, or WLS optical featureor material disposed between the respective scintillator rods.
 18. Thedetector module of claim 11, wherein the photodetector elements,including at least one of strip photodetectors with readout at both endsand continuous area photodetectors with readout at four corners, aredisposed at ends of the scintillator rods, the ends defining an end facehaving a cross sectional area transverse to the first and seconddirection in the first and second layers, respectively. 19-119.(canceled)
 120. An imaging system comprising one or more detectormodules configured to interrogate a material sample by detectingionizing radiation according to claim 1, and further comprising at leastone of an active or passive encoding technique implemented with at leastone of the detector modules or the material sample.
 121. The imagingsystem of claim 120, further comprising one or more of: an imageprocessor in communication with the detector modules, the imageprocessor configured to generate images of the material sample based onthe ionizing radiation and the encoding technique; wherein the materialsample includes biological tissue and the image processor is configuredfor medical imaging thereof; wherein the material sample includes anon-tissue material and the image processor is configured fornon-medical science, industry or inspection imaging thereof; or whereinthe encoding technique utilizes non-optical information carriers with orwithout passive optical encoding.
 122. The imaging system of claim 120,wherein the encoding technique includes modifying at least one of localor global electric, magnetic, electromagnetic, acoustic and thermalionizing radiation properties of the detector modules.
 123. The imagingsystem of claim 122, wherein one or more of: the radiation detectormodules include one or more of slab, block, pixelated, array, layered,3D structured and structured ionizing radiation detector elements; orthe encoding technique is implemented with ionizing radiation structureddetector elements of the detector modules.
 124. An imaging systemcomprising one or more detector modules configured to interrogate amaterial sample by detecting ionizing radiation according to claim 11,wherein the encoding pattern comprises at least one of active or passiveencoding on at least one of the detector modules.
 125. The imagingsystem of claim 124, further comprising an image processor incommunication with the detector modules, the image processor configuredto generate images of the material sample based on the ionizingradiation and the encoding technique, wherein: the material sampleincludes biological tissue and the image processor is configured formedical imaging thereof; or the material sample includes a non-tissuematerial and the image processor is configured for non-medical science,industry or inspection imaging thereof.
 126. The imaging system of claim124, wherein the encoding includes: modifying at least one electric,magnetic, electromagnetic, acoustic or thermal ionizing radiationproperty of the at least one detector module; non-optical informationcarriers with or without passive optical encoding; or ionizing radiationstructured detector elements of the at least one detector module. 127.The imaging system of claim 126, wherein the radiation detector modulesinclude one or more of slab, block, pixelated, array, layered, 3Dstructured and structured ionizing radiation detector elements.